Processes for making crush recoverable polymer scaffolds

ABSTRACT

Methods for making scaffolds for delivery via a balloon catheter are described. The scaffold, after being deployed by the balloon, provides a crush recovery of about 90% after the diameter of the scaffold has been pinched or crushed by 50%. The scaffold structure has patterns that include an asymmetric or symmetric closed cell, and links connecting such closed cells.

PRIORITY CLAIM

This application claims priority as a continuation of U.S. applicationSer. No. 13/090,164 filed Apr. 19, 2011, continuation-in-part of U.S.application Ser. No. 13/015,474 filed Jan. 27, 2011 and U.S. applicationSer. No. 13/015,488 filed Jan. 27, 2011. This application also claimspriority to U.S. provisional application No. 61/385,891 filed on Sep.23, 2010, and U.S. provisional application No. 61/385,902 filed Sep. 23,2010.

FIELD OF THE INVENTION

The present invention relates to drug-eluting medical devices; moreparticularly, this invention relates to polymeric scaffolds that areexpanded by a delivery balloon.

BACKGROUND OF THE INVENTION

Radially expandable endoprostheses are artificial devices adapted to beimplanted in an anatomical lumen. An “anatomical lumen” refers to acavity, duct, of a tubular organ such as a blood vessel, urinary tract,and bile duct. Stents are examples of endoprostheses that are generallycylindrical in shape and function to hold open and sometimes expand asegment of an anatomical lumen (one example of a stent is found in U.S.Pat. No. 6,066,167 to Lau et al). Stents are often used in the treatmentof atherosclerotic stenosis in blood vessels. “Stenosis” refers to anarrowing or constriction of the diameter of a bodily passage ororifice. In such treatments, stents reinforce the walls of the bloodvessel and prevent restenosis following angioplasty in the vascularsystem. “Restenosis” refers to the reoccurrence of stenosis in a bloodvessel or heart valve after it has been treated (as by balloonangioplasty, stenting, or valvuloplasty) with apparent success.

The treatment of a diseased site or lesion with a stent involves bothdelivery and deployment of the stent. “Delivery” refers to introducingand transporting the stent through an anatomical lumen to a desiredtreatment site, such as a lesion. “Deployment” corresponds to expansionof the stent within the lumen at the treatment region. Delivery anddeployment of a stent are accomplished by positioning the stent aboutone end of a catheter, inserting the end of the catheter through theskin into an anatomical lumen, advancing the catheter in the anatomicallumen to a desired treatment location, expanding the stent at thetreatment location, and removing the catheter from the lumen.

In the case of a balloon expandable stent, the stent is mounted about aballoon disposed on the catheter. Mounting the stent typically involvescompressing or crimping the stent onto the balloon prior to insertion inan anatomical lumen. At the treatment site within the lumen, the stentis expanded by inflating the balloon. The balloon may then be deflatedand the catheter withdrawn from the stent and the lumen, leaving thestent at the treatment site. In the case of a self-expanding stent, thestent may be secured to the catheter via a retractable sheath. When thestent is at the treatment site, the sheath may be withdrawn which allowsthe stent to self-expand.

The stent must be able to satisfy a number of basic, functionalrequirements. The stent must be capable of withstanding the structuralloads, for example, radial compressive forces, imposed on the stent asit supports the walls of a vessel after deployment. Therefore, a stentmust possess adequate radial yield strength. After deployment, the stentmust adequately maintain its size and shape throughout its service lifedespite the various forces that may come to bear on it. In particular,the stent must adequately maintain a vessel at a prescribed diameter fora desired treatment time despite these forces. The treatment time maycorrespond to the time required for the vessel walls to remodel, afterwhich the stent is no longer necessary for the vessel to maintain adesired diameter.

Radial yield strength, which is the ability of a stent to resist radialcompressive forces, relates to a stent's radial yield strength andradial stiffness around a circumferential direction of the stent. Astent's “radial yield strength” or “radial strength” (for purposes ofthis application) may be understood as the compressive loading, which ifexceeded, creates a yield stress condition resulting in the stentdiameter not returning to its unloaded diameter, i.e., there isirrecoverable deformation of the stent. When the radial yield strengthis exceeded the stent is expected to yield more severely and only aminimal additional force is required to cause major deformation.

Even before the radial yield strength is exceeded there may be permanentdeformation in the stent following radial compressive load, but thisdegree of permanent deformation somewhere in the stent is not severeenough to have a significant effect on the stent's overall ability toradially support a vessel. Therefore, in some cases the art may view“radial yield strength” as the maximum radial loading, beyond which thescaffold stiffness changes dramatically. “Radial yield strength” unitsare sometimes force-divided-by-stent length, which is an expression ofradial yield strength on a per-unit-length basis. Thus, for a radialyield strength per unit length, e.g., F N/mm, the radial load which, ifit exceeds this value, would result in significant change in stiffnessfor a stent having two different lengths, L1 and L2, would therefore bethe product F*L1 and F*L2, respectively. The value F, however, is thesame in both cases, so that a convenient expression can be used toappreciate the radial yield strength independent of the length of thestent. Typically, the radial force that identifies the point wherestiffness is lost does not change much on a per-unit-length basis whenthe stent length changes.

Stents implanted in coronary arteries are primarily subjected to radialloads, typically cyclic in nature, which are due to the periodiccontraction and expansion of vessels as blood is pumped to and from abeating heart. Stents implanted in peripheral blood vessels, or bloodvessels outside the coronary arteries, e.g., iliac, femoral, popliteal,renal and subclavian arteries, however, must be capable of sustainingboth radial forces and crushing or pinching loads. These stent types areimplanted in vessels that are closer to the surface of the body. Becausethese stents are close to the surface of the body, they are particularlyvulnerable to crushing or pinching loads, which can partially orcompletely collapse the stent and thereby block fluid flow in thevessel.

As compared to a coronary stent, which is designed to counteractprimarily radial loads, a peripheral stent must take into account thesignificant differences between pinching or crushing loads and radialloads, as documented in Duerig, Tolomeo, Wholey, Overview ofsuperelastic stent Design, Min Invas Ther & Allied Technol 9(3/4), pp.235-246 (2000) and Stoeckel, Pelton, Duerig, Self-Expanding NitinolStents—Material and Design Considerations, European Radiology (2003).The corresponding crushing and radial stiffness properties of the stentalso can vary dramatically. As such, a stent that possesses a certaindegree of radial stiffness does not, generally speaking, also indicatethe degree of pinching stiffness possessed by the stent. The twostiffness properties are not the same, or even similar.

The amount of cross-sectional crush during walking expected for aperipheral stent implanted within the femoral artery has been estimatedto be about 5.8+/−7%, 6.5+/−4.9% and 5.1+/−6.4% at the top, middle andbottom portions of the femoral artery in older patients and 2.5+/−7.7%,−0.8+/−9.4% and −1.5+/−10.5% for younger patients. These crush estimatescan correspond to times when the patient is walking. Significant morecrush could occur occasionally by external forces. Other considerationsfor peripheral stents are the degree of bending and axial compressionthe stent can withstand without mechanical loss of strength/stiffness.As compared to coronary stents, a peripheral stent usually has lengthsof between about 20 and 200 mm when implanted in the superficial femoralartery, as an example. As such, the stent must be flexible enough towithstand axial compression and bending loading without failure. Theamount of bending and axial compression expected has been studied andreported in Nikanorov, Alexander, M. D. et al., Assessment ofself-expanding Nitinol stent deformation after chronic implantation intothe superficial femoral artery.

To date the most commonly used type of peripheral stent areself-expanding stents made from super-elastic material, such as Nitinol.This type of material is known for its ability to return to its originalconfiguration after severe deformation, such as a crushing load orlongitudinal bending. However, this variety of self-expanding stentshave undesired qualities; most notably, the high resiliency ofsuper-elastic material produces what is commonly referred to as a“chronic outward force” (COF) on the blood vessel supported by thestent. Complications resulting from COF are discussed in Schwartz, LewisB. et al. Does Stent Placement have a learning curve: what mistakes dowe as operators have to make and how can they be avoided?, AbbottLaboratories; Abbott Park, Ill., USA. It is believed that a COF exertedon a blood vessel by a self-expending stent is a main contributor tohigh degrees of restenosis of lesions treated by the self-expandingstent. It has been shown that not even an anti-proliferative drugdelivered from drug eluting self-expandable stents can mitigate therestenosis caused by the stent's COF.

Stents that are plastically deformed by a balloon to support a vessel donot suffer from this drawback. Indeed, balloon expanded stents, incontrast to self-expanding stents made from a super-elastic material,have the desirable quality of being deployable to the desired diameterfor supporting the vessel without exerting residual outward forces onthe vessel. However, the prior art has concluded that plasticallydeformed stents, once collapsed, pinched or crushed in a peripheralartery will remain so, permanently blocking the vessel. The prior arthas concluded, therefore, that plastically deformed stents pose anundesirable condition to the patient and should not be used to treatperipheral blood vessels.

A polymer scaffold, such as that described in US 2010/0004735 is madefrom a biodegradable, bioabsorbable, bioresorbable, or bioerodablepolymer. The terms biodegradable, bioabsorbable, bioresorbable,biosoluble or bioerodable refer to the property of a material or stentto degrade, absorb, resorb, or erode away from an implant site. Thepolymer scaffold described in US 2010/0004735, as opposed to a metalstent, is intended to remain in the body for only a limited period oftime. The scaffold is made from a biodegradable or bioerodable polymer.In many treatment applications, the presence of a stent in a body may benecessary for a limited period of time until its intended function of,for example, maintaining vascular patency and/or drug delivery isaccomplished. Moreover, it is believed that biodegradable scaffoldsallow for improved healing of the anatomical lumen as compared to metalstents, which may lead to a reduced incidence of late stage thrombosis.In these cases, there is a desire to treat a vessel using a polymerscaffold, in particular a bioerodible polymer scaffold, as opposed to ametal stent, so that the prosthesis's presence in the vessel is for alimited duration. However, there are numerous challenges to overcomewhen developing a polymer scaffold.

The art recognizes a variety of factors that affect a polymericscaffold's ability to retain its structural integrity and/or shape whensubjected to external loadings, such as crimping and balloon expansionforces. These interactions are complex and the mechanisms of action notfully understood. According to the art, characteristics differentiatinga polymeric, bio-bioresorbable scaffold of the type expanded to adeployed state by plastic deformation from a similarly functioning metalscaffold are many and significant. Indeed, several of the acceptedanalytic or empirical methods/models used to predict the behavior ofmetallic scaffolds tend to be unreliable, if not inappropriate, asmethods/models for reliably and consistently predicting the highlynon-linear, time dependent behavior of a polymeric load-bearingstructure of a balloon-expandable scaffold. The models are not generallycapable of providing an acceptable degree of certainty required forpurposes of implanting the scaffold within a body, orpredicting/anticipating the empirical data.

Moreover, it is recognized that the state of the art in medicaldevice-related balloon fabrication, e.g., non-compliant balloons forscaffold deployment and/or angioplasty, provide only limited informationabout how a polymeric material might behave when used to support a lumenwithin a living being via plastic deformation of a network of ringsinterconnected by struts. In short, methods devised to improvemechanical features of an inflated, thin-walled balloon structure, mostanalogous to mechanical properties of a pre-loaded membrane when theballoon is inflated and supporting a lumen, simply provides little, ifany insight into the behavior of a deployed polymeric scaffold. Onedifference, for example, is the propensity for fracture or cracks todevelop in a polymer scaffold. The art recognizes the mechanical problemas too different to provide helpful insights, therefore, despite ashared similarity in class of material. At best, the balloon fabricationart provides only general guidance for one seeking to improvecharacteristics of a balloon-expanded, bio-absorbable polymericscaffold.

Polymer material considered for use as a polymeric scaffold, e.g.poly(L-lactide) (“PLLA”), poly(L-lactide-co-glycolide) (“PLGA”),poly(D-lactide-co-glycolide) or poly(L-lactide-co-D-lactide)(“PLLA-co-PDLA”) with less than 10% D-lactide, and PLLD/PDLA stereocomplex, may be described, through comparison with a metallic materialused to form a stent, in some of the following ways. A suitable polymerhas a low strength to weight ratio, which means more material is neededto provide an equivalent mechanical property to that of a metal.Therefore, struts must be made thicker and wider to have the requiredstrength for a scaffold to support lumen walls at a desired radius. Thescaffold made from such polymers also tends to be brittle or havelimited fracture toughness. The anisotropic and rate-dependant inelasticproperties (i.e., strength/stiffness of the material varies dependingupon the rate at which the material is deformed) inherent in thematerial, only compound this complexity in working with a polymer,particularly, bio-absorbable polymer such as PLLA or PLGA.

Processing steps performed on, and design changes made to a metal stentthat have not typically raised concerns, or required careful attentionto unanticipated changes in the average mechanical properties of thematerial, therefore, may not also apply to a polymer scaffold due to thenon-linear and sometimes unpredictable nature of the mechanicalproperties of the polymer under a similar loading condition. It issometimes the case that one needs to undertake extensive validationbefore it even becomes possible to predict more generally whether aparticular condition is due to one factor or another—e.g., was a defectthe result of one or more steps of a fabrication process, or one or moresteps in a process that takes place after scaffold fabrication, e.g.,crimping? As a consequence, a change to a fabrication process,post-fabrication process or even relatively minor changes to a scaffoldpattern design must, generally speaking, be investigated more thoroughlythan if a metallic material were used instead of the polymer. Itfollows, therefore, that when choosing among different polymericscaffold designs for improvement thereof, there are far less inferences,theories, or systematic methods of discovery available, as a tool forsteering one clear of unproductive paths, and towards more productivepaths for improvement, than when making changes in a metal stent.

The present inventors recognize, therefore, that, whereas inferencespreviously accepted in the art for stent validation or feasibility whenan isotropic and ductile metallic material was used, those inferenceswould be inappropriate for a polymeric scaffold. A change in a polymericscaffold pattern may affect not only the stiffness or lumen coverage ofthe scaffold in its deployed state supporting a lumen, but also thepropensity for fractures to develop when the scaffold is being crimpedor being deployed. This means that, in comparison to a metallic stent,there is generally no assumption that can be made as to whether achanged scaffold pattern may not produce an adverse outcome, or requirea significant change in a processing step (e.g., tube forming, lasercutting, crimping, etc.). Simply put, the highly favorable, inherentproperties of a metal (generally invariant stress/strain properties withrespect to the rate of deformation or the direction of loading, and thematerial's ductile nature), which simplify the stent fabricationprocess, allow for inferences to be more easily drawn between a changedstent pattern and/or a processing step and the ability for the stent tobe reliably manufactured with the new pattern and without defects whenimplanted within a living being.

A change in the pattern of the struts and rings of a polymeric scaffoldthat is plastically deformed, both when crimped onto, and when laterdeployed by a balloon, unfortunately, is not predictable to the same orsimilar degree as for a metal stent. Indeed, it is recognized thatunexpected problems may arise in polymer scaffold fabrication steps as aresult of a changed pattern that would not have necessitated any changesif the pattern was instead formed from a metal tube. In contrast tochanges in a metallic stent pattern, a change in polymer scaffoldpattern may necessitate other modifications in fabrication steps orpost-fabrication processing, such as crimping and sterilization.

In addition to meeting the requirements described above, it is desirablefor a scaffold to be radiopaque, or fluoroscopically visible underx-rays. Accurate placement is facilitated by real time visualization ofthe delivery of a scaffold. A cardiologist or interventional radiologistcan track the delivery catheter through the patient's vasculature andprecisely place the scaffold at the site of a lesion. This is typicallyaccomplished by fluoroscopy or similar x-ray visualization procedures.For a scaffold to be fluoroscopically visible it must be more absorptiveof x-rays than the surrounding tissue. Radiopaque materials in ascaffold may allow for its direct visualization. However, a significantshortcoming of a biodegradable polymer scaffold (and polymers generallycomposed of carbon, hydrogen, oxygen, and nitrogen) is that they areradiolucent with no radiopacity. Biodegradable polymers tend to havex-ray absorption similar to body tissue. One way of addressing thisproblem is to attach radiopaque markers to structural elements of thescaffold. A radiopaque marker can be disposed within a structuralelement in such a way that the marker is secured to the structuralelement. However, the use of stent markers on a polymeric scaffoldentails a number of challenges. One challenge relates to the difficultyof insertion of markers. These and related difficulties are discussed inUS 2007/0156230.

There is a need to develop a prosthesis for treating peripheral bloodvessels that possesses the desirable qualities of a balloon expandedstent, which does not exert residual outward forces on the vessel (as inthe case of a self-expanding stent) while, at the same time, beingsufficiently resilient to recover from a pinching or crushing load in aperipheral blood vessel, in addition to the other loading eventsexpected within a peripheral blood vessel that are not typicallyexperienced by a coronary scaffold. There is also a need to fabricatesuch a polymer scaffold so that the prosthesis also is capable ofpossessing at least a minimum radial strength and stiffness required tosupport a peripheral blood vessel; a low crossing profile; and a limitedpresence in the blood vessel. There is also a need for a scaffold thatis easily monitored during its pendency using standard imagingtechniques, and is capable of high yield production.

SUMMARY OF THE INVENTION

The invention provides processes for making a polymer scaffold suited toaddress the foregoing needs including high crush recoverability, e.g.,at least about 90-95% after a 50% crushing load. The scaffold is cutfrom a polymer tube and crimped to a balloon. Accordingly, the inventionprovides processes for making a balloon expandable, plastically deformedscaffold cut from a tube and being suitable for use as a peripheralscaffold. As such, the drawbacks of self-expanding stents can beobviated by practicing the invention.

To date the art has relied on metals or alloys for support and treatmentof peripheral blood vessels. As mentioned earlier, once a metallic stentis implanted it remains in the body permanently, which is not desired. Ascaffold made from a material that dissolves after it treats an occludedvessel, therefore, would be preferred over a metal stent. A polymer,however, is much softer than a metal. If it will serve as a replacementto metal, a new design approach is needed.

High radial force, small crimped profile and crush recovery is needed inthe polymer scaffold. If the material cannot be modified enough to meetthese needs, then a modification to the design of the scaffold networkof struts is required. There are a few known approaches to increase theradial yield strength. One is to increase the wall thickness and anotheris to increase the strut width. Both of these modifications, however,will result in greater profile of the device at the crimped state. Asmall crimped profile of the device and increased stiffness and strengthis therefore necessary and heretofore has not been addressed in the art.

As will be appreciated, aspects of a polymer scaffold produced by theprocesses described herein contradict conclusions that have beenpreviously made in the art concerning the suitability of aballoon-expandable stent, or scaffold for use in peripheral bloodvessels. The problems concerning self-expanding stents are known.Therefore a replacement is sought. However, the conventional wisdom isthat a balloon expanded stent having sufficient radial strength andstiffness, as opposed to a self-expanding stent, is not a suitablereplacement, especially in vessels that will impose high bending and/orcrushing forces on the implanted prosthesis.

According to the invention, crush-recoverable polymer scaffolds made bythe processes described herein possess a desired radial stiffness andyield strength, fracture toughness and capability of being crimped downto a target delivery diameter that will properly balance three competingdesign attributes: radial yield strength and stiffness versus fracturetoughness, in-vivo performance versus compactness for delivery to avessel site, and crush recovery versus radial yield strength andstiffness.

Disclosed herein are embodiments of methods for making a scaffold thatcan effectively balance these competing needs, thereby providing analternative to prostheses that suffer from chronic outward force. Aswill be appreciated from the disclosure, various polymer scaffoldcombinations were fabricated from the inventive methods and tested inorder to better understand the characteristics of a scaffold that mightaddress at least the following needs: crush recoverability of thescaffold without sacrificing a desired minimal radial stiffness andstrength, recoil, deploy-ability and crimping profile; acute recoil atdeployment—the amount of diameter reduction within ½ hour of deploymentby the balloon; delivery/deployed profile—i.e., the amount the scaffoldcould be reduced in size during crimping while maintaining structuralintegrity; in vitro radial yield strength and radial stiffness; minimizecrack formation/propagation/fracture when crimped and expanded by theballoon, or when implanted within a vessel and subjected to acombination of bending, axial compression, radial crush and radialcompressive loads; uniformity of deployment of scaffold rings whenexpanded by the balloon; and adequate pinching/crushing stiffness.

The various attributes of a crush-recoverable scaffold, such as thevarious ratios of scaffold properties, e.g., strut and link dimensions(as discussed herein) and relationships relating to crushrecoverability; acute recoil at deployment; delivery/deployed profile;in vitro radial yield strength and radial stiffness; radial yieldstrength and stiffness; uniformity of deployment of scaffold rings whenexpanded by the balloon; and pinching/crushing stiffness that wereproduced according to the methods described herein are also described inrelated U.S. application Ser. No. 13/015,474 and Ser. No. 13/015,488,which is considered part of this disclosure.

In one aspect of the invention, a crush recoverable scaffold was crimpedfrom a 7 mm, 8 mm and 9 mm outer diameter to a 2 mm outer diameter anddeployed without fracture and/or excessive cracking of struts that are atypical concern when a polymer, especially a brittle polymer like PLLA,is used to form the scaffold structure.

In another aspect of the invention a symmetric, closed cell for ascaffold improves deployment uniformity and reduces fracture problemsfor a scaffold having crush recoverability.

In another aspect of invention a method for making a medical deviceincluding a scaffold crimped to a balloon-catheter includes the steps ofbiaxially expanding a polymer precursor to form an expanded tube,forming a scaffold from the expanded tube using a laser, crimping thescaffold to a balloon-catheter at a crimping temperature between about 1to 10 degrees less than the glass transition temperature of the polymermaterial and fitting a removable sheath over the scaffold immediatelyfollowing crimping to limit recoil of the crimped scaffold, wherein thecrimped scaffold, when deployed, is capable of regaining at least 90% ofits diameter after being crushed to at least 75% of its diameter (orcrushed by an amount equal to at least about 25% of its diameter).

In another aspect of invention a method for making balloon-expandablemedical device for being implanted in a peripheral vessel of the bodyincludes the steps of biaxially expanding a polymer precursor to from anexpanded tube, and forming a scaffold from the expanded tube, includingforming struts joined at crowns to form rings, a zero angle radius atthe crowns, and symmetric closed cells formed by the rings andconnecting links connecting the rings, wherein the scaffold is capableof regaining more than 90% of its diameter after being crushed to about50% of its diameter.

In another aspect of invention a method for crimping aballoon-expandable medical device includes the steps of forming apre-crimp scaffold, including the steps of biaxially expanding a PLLAprecursor to form an expanded tube including the steps of applying apressure of about 100-120 psi, a heating nozzle rate of about 0.5-0.9mm/sec and a temperature of about 230-240 Deg Fahrenheit, and using apico-second laser, forming a scaffold from the expanded tube, includingforming struts forming ring structures connected by longitudinal links,where there are no more than four links connecting adjacent rings; andmounting the scaffold to a balloon, including the steps of using acrimping temperature between 5 to 15 degrees less than the glasstransition temperature of PLLA, and maintaining an inflated deliveryballoon during crimping to support the scaffold during crimping; whereinthe scaffold is capable of regaining more than 90% of its diameter afterbeing crushed to about 50% of its diameter.

INCORPORATION BY REFERENCE

All publications and patent applications mentioned in this specificationare herein incorporated by reference to the same extent as if eachindividual publication or patent application was specifically andindividually indicated to be incorporated by reference, and as if eachsaid individual publication or patent application was fully set forth,including any figures, herein.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1 is a perspective view of a deformed polymer tube. The tube isformed into a scaffold.

FIGS. 2A-2D are schematic drawings describing a process for radial andaxial expansion of a precursor that is formed into the tube of FIG. 1.

FIG. 3 is a partial planar view of a scaffold pattern according to afirst embodiment of a scaffold.

FIG. 4 is a partial perspective view of a scaffold structure.

FIG. 5 is a partial planar view of a scaffold pattern according to asecond embodiment of a scaffold.

FIG. 6A is a planar view of a portion of the scaffold pattern of FIG. 5taken at section VA-VA.

FIG. 6B is a planar view of a portion of the scaffold pattern of FIG. 3taken at section VB-VB.

FIGS. 7A and 7B are tables showing examples of scaffold features inaccordance with aspects of the disclosure.

FIG. 8a -8B shows a scaffold crown formation in its expanded and crimpedstates.

FIG. 8C-8D shows a scaffold crown formation in its expanded and crimpedstates for a scaffold according to the first embodiment.

FIG. 8E-8F shows a scaffold crown formation in its expanded and crimpedstates for a scaffold according to an alternative embodiment.

FIGS. 9B, 9C and 9D are scanning electron microscope (SEM) photographsof scaffold crowns. The crowns have an inner radius of about 0.00025inches. The photographs are taken after the scaffold was expanded by aballoon.

FIGS. 9A, 9F and 9G are scanning electron microscope (SEM) photographsof scaffold crowns having an inner radius substantially higher than theinner radius of the scaffold crowns in FIGS. 9B, 9C and 9D. Thephotographs are taken after the scaffold was expanded by a balloon.

FIGS. 10A-10B show the first embodiment of a scaffold including aradiopaque marker structure formed on a link connecting rings. FIG. 10Ashows the expanded configuration and FIG. 10B shows the location of theradiopaque markers relative to folded struts of the scaffold rings inthe crimped configuration.

FIGS. 11A, 11B and 11C are diagrams describing a relationship betweencrush recoverability and wall thickness for a scaffold. FIG. 11A shows across-section of a scaffold in its un-deformed (unloaded) state anddeformed state when subjected to a pinching load (drawn in phantom).FIGS. 11B-11C are models of equivalent half-cylinder shells of differentthickness to show the effects of wall thickness on crush-recoverabilitywhen a scaffold is subject a pinching load. The spring elementsconnected to points A and B are included in FIGS. 11B-11C for purposesof illustrating, by way of analogy, the change in % strain energy at theends of the half-shells (where failure occurs when the scaffold ispinched beyond its recovery point) as compared to the shells themselvesas the wall thickness increases.

FIG. 12A shows a scaffold-catheter assembly with a sheath disposed overa scaffold crimped to a balloon. FIGS. 12B-12C show side and frontalviews of the sheath of FIG. 12A.

DETAILED DESCRIPTION OF EMBODIMENTS

The disclosure proceeds as follows. First, definitions of terms that maybe used during the course of the subsequent disclosure are explained.Next, the methods for making the scaffold are explained with reference,primarily, to two embodiments of a scaffold (see FIGS. 3-7). Thedisclosure further provides results from testing of scaffold samplesmade according to the disclosure, either herein or by way ofincorporation by reference with related applications.

An expansion process for making an expanded polymer tube from anextruded polymer precursor is described, followed by a laser cuttingprocess for forming the scaffold backbone from the expanded tube.Examples of patterns are illustrated in FIGS. 3 and 5. Placement ofradiopaque markers is described next, followed by methods for applying adrug-polymer coating to the scaffold. Finally, a crimping process isdescribed. Following the above description, the disclosure explainsdesirable attributes of scaffolds produced by one or more of the abovemethods. Finally, test results are discussed including results forscaffold yield strength, stiffness, and crush recovery.

For purposes of this disclosure, the following terms and definitionsapply:

“Inflated diameter” or “expanded diameter” refers to the maximumdiameter the scaffold attains when its supporting balloon is inflated toexpand the scaffold from its crimped configuration to implant thescaffold within a vessel. The inflated diameter may refer to apost-dilation diameter which is beyond the nominal balloon diameter,e.g., a 6.5 mm semi-compliant PEBAX balloon has about a 7.4 mmpost-dilation diameter. The scaffold diameter, after attaining itsinflated diameter by balloon pressure, will to some degree decrease indiameter due to recoil effects and/or compressive forces imposed by thewall of the vessel after the balloon is removed. For instance, referringto an expansion of the V59 scaffold having the properties in FIG. 6B,when placed on a 6.5 mm PEBAX balloon and the balloon is expanded to apost-dilation condition outside a vessel, the scaffold inner diameterwill be about 7.4 mm and about (0.955)×(7.4 mm) before and after,respectively, acute-recoil has occurred. The inflated diameter may beabout 1.2 times the average vessel diameter and peripheral vessel sizestypically range from about 4 to 10 mm for purposes of this disclosure.

“Theoretical minimum diameter” means the smallest diameter for ascaffold based on its geometry of strut lengths thickness and widths. A“theoretical minimum diameter” is not defined in terms of a minimumcrimped profile for a scaffold or stent that can be later deployed andwork properly as a balloon-expanded prosthesis. Rather, it is only adefinition defined by the geometry, or minimum volume of space that adevice can occupy following a uniform reduction in diameter. As aformula, the “theoretical minimum diameter” (Dmin) may be expressed asfollows:Dmin=(Σ Swi+Σ Crj+Σ Lwk)*(1/π)+2*WT  (EQ. 3)Where the quantities above are taken from a cross-sectional slice of thescaffold,

-   Σ Swi (i=1. . . n) is the sum of n ring struts having width Swi;-   Σ Crj (j=1. . . m) is the sum of m crown inner radii having radii    Crj (times 2);-   Σ Lwk (k=1. . . p) is the sum of p links having width Lwk; and    -   WT is the scaffold wall thickness.

EQ. 3 assumes the width for a folded pair of struts, e.g., struts 420,422 in FIG. 8A is the same whether measured near the crown 410 or thestrut mid width. When the crown is built up more, so that the width iswider there than ring strut mid-width, Swi would be measured by thewidth at the crown. Also, the minimum space between struts is defined bytwice the inner radius of the adjacent crown (or valley), i.e., Crj.

For the scaffold dimensions of FIG. 7B the crown width is wider than thestrut mid-width. Therefore, using EQ. 3 Dmin is[16*(0.013)+12*(0.0005)+4*(0.0115)]*(1/π)+2*(0.011)=0.1048 in or 2.662mm (minimum diameter computed at cross-section passing through crowns).If, instead the cross-section were taken at the strut mid width (0.0116instead of 0.013) EQ. 3 gives 0.0976 in or 2.479 mm.

It should be noted that EQ. 3 assumes the struts have essentially asquare cross-section. This is the case for the scaffold of FIG. 7B(strut cross-sectional dimension at the crown is 0.011×0.013). For ascaffold having struts with a trapezoidal cross section, e.g., ascaffold cut from a smaller diameter so that the ratio of wall thicknessto outer diameter is much higher than in the case of FIG. 1, a moreaccurate approximation for Dmin would be (Σ Swi+Σ Crj+Σ Lwk)*(1/π) sincethe edges of the struts at the outer surface would abut at Dmin beforethe surfaces extending over the thickness of a strut abut each other.

The glass transition temperature (referred to herein as “T_(g)”) is thetemperature at which the amorphous domains of a polymer change from abrittle vitreous state to a solid deformable or ductile state atatmospheric pressure. In other words, T_(g) corresponds to thetemperature where the onset of segmental motion in the chains of thepolymer occurs. T_(g) of a given polymer can be dependent on the heatingrate and can be influenced by the thermal history of the polymer.Furthermore, the chemical structure of the polymer heavily influencesthe glass transition by affecting mobility of polymer chains.

“Stress” refers to force per unit area, as in the force acting through asmall area within a plane within a subject material. Stress can bedivided into components, normal and parallel to the plane, called normalstress and shear stress, respectively. Tensile stress, for example, is anormal component of stress that leads to expansion (increase in length)of the subject material. In addition, compressive stress is a normalcomponent of stress resulting in compaction (decrease in length) of thesubject material.

“Strain” refers to the amount of expansion or compression that occurs ina material at a given stress or load. Strain may be expressed as afraction or percentage of the original length, i.e., the change inlength divided by the original length. Strain, therefore, is positivefor expansion and negative for compression.

“Modulus” may be defined as the ratio of a component of stress or forceper unit area applied to a material divided by the strain along an axisof applied force that result from the applied force. For example, amaterial has both a tensile and a compressive modulus.

“Toughness”, or “fracture toughness” is the amount of energy absorbedprior to fracture, or equivalently, the amount of work required tofracture a material. One measure of toughness is the area under astress-strain curve from zero strain to the strain at fracture. Thestress is proportional to the tensile force on the material and thestrain is proportional to its length. The area under the curve then isproportional to the integral of the force over the distance the polymerstretches before breaking. This integral is the work (energy) requiredto break the sample. The toughness is a measure of the energy a samplecan absorb before it breaks. There is a difference between toughness andstrength. A material that is strong, but not tough is said to bebrittle. Brittle materials are strong, but cannot deform very muchbefore breaking.

As used herein, the terms “axial” and “longitudinal” are usedinterchangeably and refer to a direction, orientation, or line that isparallel or substantially parallel to the central axis of a scaffold orthe central axis of a tubular construct. The term “circumferential”refers to the direction along a circumference of the scaffold or tubularconstruct. The term “radial” refers to a direction, orientation, or linethat is perpendicular or substantially perpendicular to the central axisof the scaffold or the central axis of a tubular construct and issometimes used to describe a circumferential property, i.e. radial yieldstrength.

The term “crush recovery” is used to describe how the scaffold recoversfrom a pinch or crush load, while the term “crush resistance” is used todescribe the pinch forces required to cause a permanent deformation of ascaffold. A scaffold that does not possess good crush recovery does notsubstantially return to its original diameter following removal of acrushing force. As noted earlier, a scaffold having a desired radialforce can have an unacceptable crush recovery. And a scaffold having adesired crush recovery can have an unacceptable radial force.

The polymer scaffold illustrated in FIG. 4 is formed from apoly(L-lactide) (“PLLA”) tube 101 as depicted in FIG. 1. The process forforming tube 101 begins with extrusion of a tube precursor. Raw PLLAresin material heated above the melt temperature of the polymer is thenextruded through a die at a preferred extrusion temperature of 450 Deg.Fahrenheit. Further details of this step of the process are described inUS 2011/0066222 (hereinafter the '222 publ.). A carefully controlledradial and axial expansion of the formed precursor, preferably using aform of blow-molding, follows. This expansion process is employed toproduce desired mechanical properties of the scaffold, starting with theprecursor. The desirable properties include dimensional andmorphological uniformity, e.g., crystallinity, wall thickness and“roundness”, yield strength, stiffness and fracture toughness. Theresulting tube 101 is then formed into a scaffold using a laser cuttingprocess.

Expansion of the precursor is undertaken using carefully controlledparameters including pressure, rate and temperature during the expansionof the precursor. Expansion preferably occurs in both the axial andradial direction by prescribed amounts to achieve desired results. ThePLLA precursor is heated above the PLLA glass transition temperature(i.e., 60-70 degrees C.) but below the melt temperature (165-175 degreesC.), preferably, around 110-120 degrees C.

The preferred blow molding process deforms the precursor progressivelyat a predetermined longitudinal speed along the longitudinal axis of theprecursor. The tube deformation process is intended to orient polymerchains in radial and/or biaxial directions. As mentioned above, theorientation or deformation causing re-alignment is performed accordingto a precise selection of processing parameters, e.g. pressure, heat(i.e., temperature), deformation rate, to affect material crystallinityand type of crystalline formation during the deformation process.

In an alternative embodiment the tube may be made ofpoly(L-lactide-co-glycolide), poly(D-lactide-co-glycolide) (“PLGA”),polycaprolactone, (“PCL”), and other suitable semi-crystallinecopolymers or blends of these polymers. Rubber toughen material couldalso be achieved by using block copolymers or polymer blends of theabove materials in combination with low T_(g) materials such aspolycprolactone, polyethyleneglycol and polydioxanone. Alternativelymultilayered structures could be extruded. Material choices may belimited when taking into account the complex loading environmentassociated with many peripheral vessel locations, particularly thoselocated close to limbs.

The degree of radial expansion that the polymer tube undergoes canpartially characterize the degree of induced circumferential molecularand crystal orientation as well as strength in a circumferentialdirection. The degree of radial expansion is quantified by a radialexpansion (“RE”) ratio, defined as RE Ratio=(Inside Diameter of ExpandedTube)/(Original Inside Diameter of the tube). The RE ratio can also beexpressed as a percentage, defined as RE %=(RE ratio−1).times.100%. Thedegree of axial extension that the polymer tube undergoes can partiallycharacterize induced axial molecular or crystal orientation as well asstrength in an axial direction. The degree of axial extension isquantified by an axial extension (“AE”) ratio, defined as AERatio=(Length of Extended Tube)/(Original Length of the Tube). The AEratio can also be expressed as a percentage, defined as AE %=(AEratio−1).times.100%.

Blow molding includes first positioning the tube precursor (orprecursor) in a hollow cylindrical member or mold. The mold controls thedegree of radial deformation of the precursor by limiting thedeformation of the outside diameter or surface of the precursor to theinside diameter of the mold. While in the mold, the precursortemperature is above T_(g) of PLLA to facilitate deformation. Thistemperature is a processing parameter referred to as the “expansiontemperature” or “process temperature.” The heating to the expansiontemperature can be achieved by heating a gas to the expansiontemperature and discharging the heated gas onto an exterior surface ofthe mold containing the precursor.

While in the mold, one end of the precursor is sealed or blocked. Thus,introduction of gas into the opposite end of the precursor will increaseinternal fluid pressure relative to ambient pressure in a region betweenthe outer surface of the precursor and the inner surface of the mold.The internal fluid pressure is a processing parameter referred to as the“expansion pressure” or “process pressure.” Examples of gas that may beused to create the expansion pressure include without limitation ambientair, substantially pure oxygen, substantially pure nitrogen, and othersubstantially pure inert gases. In combination with other blow moldingprocess parameters, the expansion pressure affects the rate at which theprecursor deforms radially and axially to produce the tube 101 shown inFIG. 1. Blow molding may include pulling one end of the precursor. Atensile force, which is another processing parameter, is applied to oneend of the while holding the other end of the precursor stationary.

The radially and axially deformed precursor may then be cooled fromabove T_(g) to below T_(g), either before or after decreasing thepressure and/or decreasing tension. Cooling at a controlled temperature,or rate of temperature drop, helps insure the tube 101 maintains theproper shape, size, and length following radial expansion and axialextension. Slow cooling through a temperature range between T_(m) andT_(g) might result in a loss of amorphous chain orientation and cause adecrease in fracture toughness of the finished scaffold. Preferably,though not necessarily, the deformed precursor can be cooled quickly orquenched in relatively cold gas or liquid to a temperature below T_(g)to maintain chain orientation that was formed during tubing expansion.The deformed precursor after cooling produces the tube 101, which may bethen cut to produce the scaffold described in FIGS. 3, 4, 6B and 7B.

FIGS. 2A-2D schematically depicts a molding system 500 for simultaneousradial and axial deformation of a polymer tube. FIG. 2A depicts an axialcross-section of a polymer tube 501 with an undeformed outside diameter505 positioned within a mold 510. The mold 510 limits the radialdeformation of the polymer tube 501 to a diameter 515 corresponding tothe inside diameter of the mold 510. The polymer tube 501 is closed at adistal end 520. A gas is conveyed, as indicated by an arrow 525, into anopen end 521 of the polymer tube 501 to increase internal fluid pressurewithin tube 501.

A tensile force 522 is applied to the distal end 520 in an axialdirection. In other embodiments, a tensile force is applied at theproximal end 521 and the distal end 520.

A circular band or segment of the polymer tube 500 is heated by a nozzle530. The nozzle has fluid ports that direct a heated fluid, such as hotair, at two circumferential locations of the mold 510, as shown byarrows 535 and 540. FIG. 2B depicts a radial cross-section showing thetube 501 within the mold 510, and the nozzle 530 supported by structuralmembers 560. Additional fluid ports can be positioned at othercircumferential locations of the mold 510 to facilitate uniform heatingaround a circumference of the mold 510 and the tube 501. The heatedfluid flows around the mold 510, as shown by arrows 555, to heat themold 510 and the tube 501 to a predetermined temperature above ambienttemperature.

The nozzle 530 translates along the longitudinal axis 573 of the mold510 as shown by arrows 565 and 567. That is, the nozzle 530 moveslinearly in a direction parallel to the longitudinal axis 573 of themold 510. As the nozzle 530 translates along the axis of the mold 510,the tube 501 radially deforms. The combination of elevated temperatureof the tube 501, the applied axial tension, and the applied internalpressure cause simultaneous axial and radial deformation of the tube501, as depicted in FIGS. 2C and 2D.

FIG. 2C depicts the system 500 with an undeformed section 571, adeforming section 572, and a deformed section 570 of the polymer tube501. Each section 570, 571, 572 is circular in the sense that eachsection extends completely around the central axis 573. The deformingsection 572 is in the process of deforming in a radial direction, asshown by arrow 580, and in an axial direction, as shown by arrow 582.The deformed section 570 has already been deformed and has an outsidediameter that is the same as the inside diameter of the mold 510.

FIG. 2D depicts the system 500 at some time period after FIG. 2C. Thedeforming section 572 in FIG. 2D is located over a portion of what wasan undeformed section in FIG. 2C. Also, the deformed section 570 in FIG.2D is located over what was the deforming section 572 in FIG. 2C. Thusit will be appreciated that the deforming section 572 propagateslinearly along the longitudinal axis 573 in the same general direction565, 567 that the heat sources 530 are moving.

In FIG. 2D, the deforming section 572 has propagated or shifted by anaxial distance 574 from its former position in FIG. 2C. The deformedsection 570 has grown longer by the same axial distance 574. Deformationof the tube 501 occurs progressively at a selected longitudinal ratealong the longitudinal axis 573 of the tube. Also, the tube 501 hasincreased in length by a distance 523 compared to FIG. 2C.

Depending on other processing parameters, the speed at which the heatsources or nozzles 530 are linearly translated over the mold 510 maycorrespond to the longitudinal rate of propagation (also referred to asthe axial propagation rate) of the polymer tube 501. Thus, the distance574 that the heat sources 530 have moved is the same distance 575 thatthe deformed section 570 has lengthened.

The rate or speed at which the nozzles 530 are linearly translated overthe mold 510 is a processing parameter that relates to the amount oftime a segment of the polymer tube is heated at the expansiontemperature and the uniformity of such heating in the polymer tubesegment.

It is to be understood that the tensile force, expansion temperature,and expansion pressure are applied simultaneously to the tube 501 whilethe nozzle 530 moves linearly at a constant speed over the mold. Again,the “expansion pressure” is the internal fluid pressure in the polymertube while it is blow molded inside the mold. In FIGS. 2A-2D, the“expansion temperature” is the temperature to which a limited segment ofthe polymer tube is heated during blow molding. The “limited segment” isthe segment of the polymer tube surrounded by the nozzle 530. The“limited segment” may include the deforming section 572. The heating ofthe polymer tube to the expansion temperature can be achieved by heatinga gas to the expansion temperature and discharging the heated gas fromthe nozzle 530 onto the mold 510 containing the polymer tube.

The processing parameters of the above-described blow molding processinclude without limitation the tensile force, expansion temperature, theexpansion pressure, and nozzle translation rate or linear movementspeed. It is expected that the rate at which the tube deforms duringblow molding depends at least upon these parameters. The deformationrate has both a radial component, indicated by arrow 580 in FIGS. 2C and2D, and an axial component, indicated by an arrow 582. It is believedthat the radial deformation rate has a greater dependence on theexpansion pressure and the axial component has a greater dependence onthe translation rate of the heat source along the axis of the tube. Itis also expected that the deformation rate is dependant upon thepre-existing morphology of the polymer in the undeformed section 571.Also, since deformation rate is a time dependent process, it is expectedto have an effect on the resulting polymer morphology of the deformedtube after blow molding.

The term “morphology” refers to the microstructure of the polymer whichmaybe characterized, at least in part, by the percent crystallinity ofthe polymer, the relative size of crystals in the polymer, the degree ofuniformity in spatial distribution of crystals in the polymer, and thedegree of long range order or preferred orientation of molecules and/orcrystals. Morphology may also refer to the degree of phase separation ina rubber-toughened material. The crystallinity percentage refers to theproportion of crystalline regions to amorphous regions in the polymer.Polymer crystals can vary in size and are sometimes geometricallyarranged around a nucleus, and such arrangement may be with or without apreferred directional orientation. A polymer crystal may grow outwardlyfrom the nucleus as additional polymer molecules join the orderedarrangement of polymer molecule chains. Such growth may occur along apreferred directional orientation.

Applicant believes that all the above-described processing parametersaffect the morphology of the deformed polymer tube 501. As used herein,“deformed tube 501” and “blow molded tube 501” are used interchangeablyand refer to the deformed section 570 of the polymer tube 501 of FIGS.2C and 2D. Without being limited to a particular theory, Applicantbelieves that increasing the crystallinity percentage will increase thestrength of the polymer but also tends to make the polymer brittle andprone to fracture when the crystallinity percentage reaches a certainlevel. Without being limited to a particular theory, Applicant believesthat having a polymer with relatively small crystal size has higherfracture toughness or resistance to fracture. Applicant also believesthat having a deformed tube 501 with spatial uniformity in the radialdirection, axial direction, and circumferential direction also improvesstrength and fracture toughness of the stent made from the deformedtube.

It should be noted that the above-described processing parameters areinterdependent or coupled to each other. That is, selection of aparticular level for one processing parameter affects selection ofappropriate levels for the other processing parameters that would resultin a combination of radial expansion, axial extension, and polymermorphology that produces a stent with improved functionalcharacteristics such as reduced incidence of strut fractures and reducedrecoil. For example, a change in expansion temperature may also changethe expansion pressure and nozzle translation rate required to obtainimproved stent functionality.

Expansion temperature affects the ability of the polymer to deform(radially and axially) while simultaneously influencing crystalnucleation rate and crystal growth rate, as shown in FIG. 4 of '222publ. It depicts an exemplary schematic plot of crystallization underquiescent condition, showing crystal nucleation rate (“RN”) and thecrystal growth rate (“RCG”) as a function of temperature. The crystalnucleation rate is the rate at which new crystals are formed and thecrystal growth rate is the rate of growth of formed crystals. Theexemplary curves for RN and RCG in FIG. 4 of '222 publ. have a curvedbell-type shape that is similar to RN and RCG curves for PLLA. Theoverall rate of quiescent crystallization (“RCO”) is the sum of curvesRN and RCG.

Quiescent crystallization can occur from a polymer melt, which is to bedistinguished from crystallization that occurs solely due to polymerdeformation. In general, quiescent crystallization tends to occur in asemi-crystalline polymer at temperatures between T_(g) and T_(m) of thepolymer. The rate of quiescent crystallization in this range varies withtemperature. Near T_(g), nucleation rate is relatively high andquiescent crystal growth rate is relatively low; thus, the polymer willtend to form small crystals at these temperatures. Near T_(m),nucleation rate is relatively low and quiescent crystal growth rate isrelatively high; thus, the polymer will form large crystals at thesetemperatures.

As previously indicated, crystallization also occurs due to deformationof the polymer. Deformation stretches long polymer chains and sometimesresults in fibrous crystals generally oriented in a particulardirection. Deforming a polymer tube made of PLLA by blow molding at aparticular expansion temperature above T_(g) results in a combination ofdeformation-induced crystallization and temperature-inducecrystallization.

As indicated above, the ability of the polymer to deform is dependent onthe blow molding temperature (“expansion temperature”) as well as beingdependant on the applied internal pressure (“expansion pressure”) andtensile force. As temperature increases above T_(g), molecularorientation is more easily induced with applied stress. Also, astemperature approaches T_(m), quiescent crystal growth rate increasesand quiescent nucleation rate decreases. Thus, it will also beappreciated that the above described blow molding process involvescomplex interaction of the processing parameters all of whichsimultaneously affect crystallinity percentage, crystal size, uniformityof crystal distribution, and preferred molecular or crystal orientation.

As mentioned earlier, in a preferred embodiment the PLLA tube was madeentirely of PLLA. The preferred levels are given below for the blowmolding process parameters for a PLLA precursor having an initial(before blow molding) crystallinity percentage of up to about 20% andmore narrowly from about 5% to about 15%. Applicants believe the blowmolding process parameter levels given below result in a deformed PLLAtube having a crystallinity percentage below 50% and more narrowly fromabout 30% to about 40%.

Approximate values of the processing conditions for a preferredembodiment are depicted below. The processing conditions provided TABLE1 were used to produce tube 101 from the precursor. The tube 101 wasthen formed into the scaffold described in FIGS. 3-7.

TABLE 1 Preferred processing conditions for producing tube 101 Expansionparameters used to form a tube 101 from an Processing parametersextruded PLLA precursor temperature (Deg. F.) 230-240 pressure (psi)100-120 nozzle rate (mm/sec) 0.48-0.88 % RA 400 % RE 40-50 Degree ofCrystallinity   40-50%

After expansion, the tube 101 may be subjected to an extended period ofelevated temperature. In one embodiment, a PLLA tube 101 is subjected toa temperature of between about 40-50 Deg Celsius or about 47 Deg Celsiusbefore laser-cutting the tube to form the scaffold. This step wouldoccur after the expanded tube is quenched. The subsequent, prolongedexposure to an elevated temperature, which may be included in theprocess, is intended to induce relaxation of internal stresses in thedeformed precursor far more slowly than a typical annealing process. Theprocess may be thought of as a “cold crystallization process”.

In a preferred embodiment a PLLA tube (or scaffold after laser cuttingof the tube) is subjected to a temperature of about 47 Deg over a fourweek period. It is believed that the process produces enhanced strengthand stiffness properties. Other temperature ranges may be used.Conceivably, a higher temperature will bring about the annealing processmore quickly. However, it is preferred to not raise the temperature toofar above about 50 Deg Celsius as this will start to cause excessive andundesired movement or re-ordering of polymer chains. For PLLA it wasfound that cold sterilization at 37 Deg Celsius did not produce anynoticeable differences.

Following expansion the tube 101 is fabricated into a scaffold by lasermachining. Material is removed from selected regions of the tube whichresults in formation of the patterns depicted herein. A laser beam isscanned over the surface of the tube 101 resulting in removal of atrench or kerf extending all the way through a wall of the tubing,either in one pass or two passes with the laser. In a presentlypreferred embodiment of a scaffold having an average wall thickness ofabout 0.011 inches two passes are made to form the scaffold pattern.After the second pass, the starting and ending points of where thelaser's kerf meet, the region surrounded by the kerf drops out or isremoved by an assisting gas. A preferred set of laser parameters for thescaffold depicted in FIGS. 3-7 is provided in TABLE 2, below. Otherdetails are found in U.S. application Ser. No. 12/797,950.

TABLE 2 Laser Machining Parameters for a crush recoverable polymerscaffold having a wall thickness of between about .008 in and .014 in:laser type 515 nm (Trumpf, green light) Tube length (mm) 80-250 wallthickness (in) .008-.014  Scaffold length (mm) 30-120 No. of passes tocut 1-4  Cutting speed (in/min) 4-10 Tube outer diameter (mm) 6-10 Laserspot/beam size (μm) 10-20  Laser rep rate (kHz) 5-50 average power (W).7-2.0 Helium gas flow (scfh) 10-30 

In one embodiment following laser machining a structure having aplurality of struts 230 and links 234 forming a pattern 200 as shown inFIG. 3 (pattern 200 is illustrated in a planar or flattened view) isformed. the pattern shown in FIG. 3 is about that of the scaffold beforecrimping and after the scaffold is plastically, or irreversibly deformedfrom its crimped state to its deployed state within a vessel by balloonexpansion. The pattern 200 of FIG. 3, therefore, represents a tubularscaffold structure (as partially shown in three dimensional space inFIG. 4), so that an axis A-A is parallel to the central or longitudinalaxis of the scaffold. FIG. 4 shows the scaffold in a state prior tocrimping or after deployment. As can be seen from FIG. 4, the scaffoldcomprises an open framework of struts and links that define a generallytubular body.

In a preferred embodiment laser machining includes the step of formingregions in connecting links to hold a radiopaque marker bead, e.g., aplatinum bead. The structure for holding marker beads in a preferredembodiment is described in FIGS. 10A-10B and the accompanying text,infra.

The size of the marker bead is proportional to the visibility of thebead when the scaffold is implanted within the body. A marker bead,therefore, is preferably sufficiently large enough so that it can easilybe seen using a fluoroscope. A process for placing a radiopaque markerbead in each of the pair of depots 500 in FIG. 10A includes thefollowing steps. A bead is held by a vacuum pick and placed over thedepot 500. Then a hand-held mandrel is used to push the bead into thechannel formed by the depot. For the scaffold “V59”, for example (FIG.7B), the diameter of the bead (e.g., about 0.009 inches) is less thanthe wall thickness of the scaffold. As such, the bead may be placed sothat it does not protrude beyond the walls of the scaffold. This allowsthe bead to be easily placed within the depot without the additionalstep of deforming the malleable bead (so that a flush outer surface isformed with adjacent luminal and abluminal walls of the scaffold). Amarker bead process according to one embodiment may therefore includeplacing the bead within the depot without significantly deforming thebead. The bead may then be secured in place by a coating material, e.g.,a drug-polymer layer applied during a coating step for the scaffold.

Following placement of the marker beads in the scaffold, a drug-polymercoating is applied by a spraying and drying process. A spray nozzle isused to apply the coating material. And a dryer is used to applyinter-pass drying of coating. The term “inter-pass drying” means drying,or removing solvent between one, two, three or more spray passes.Preferably between about 13-14 coats are applied to reach 100% of acoating weight for a scaffold. Methods for spraying and drying ascaffold are described in U.S. application Ser. No. 12/554,820. Anapparatus and process for removing gases from a solution sprayed onto ascaffold are described in U.S. application Ser. No. 13/039,192. Duringthe spraying step, the scaffold is held on a rotating mandrel designedto minimize the amount of coating defects on the scaffold. Examples ofsuch a mandrel are described in U.S. application Ser. No. 12/752,983.

After the desired coating has been applied and the desired amount ofsolvent removed, the scaffold is mounted onto a balloon catheter by acrimping process. During crimping there is a precise control oftemperature within a specific range for these materials, in relation totheir glass transition temperature (T_(g)), to improve the retentionforce on a balloon without causing adverse effects on the polymerscaffold's yield strength and stiffness properties when it is laterexpanded by the balloon. Additionally, crimping at below but close toT_(g) has been found to reduce the instances of cracking of scaffoldstruts. Processes for crimping a scaffold are described in U.S.application Ser. No. 12/772,116 and U.S. application Ser. No.12/861,719.

Crimping processes for the “V23” scaffold (FIG. 7A) and “V59” scaffold(FIG. 7B) are summarized in TABLES 3A and 3B. In these example acrimping temperature of between 5-15 degrees below T_(g), and morepreferably about 44-52 and 48 Deg Celsius was used.

TABLE 3A crimping process for the “V23” scaffold Stage 1 - crimp headcloses to .314 in at a speed of .5 inches per second (in/s) thenimmediately go to Stage 2. Stage 2 - crimp head closes to .300 in at aspeed of .005 in/s and dwells for 30 seconds. Stage 3 - crimp headcloses to .270 in at a speed of .005 in/s and dwells for 30 seconds.Turn stopcock to release pressure from the inflated support ballooncatheter. Stage 4 - crimp head closes to .240 in at a speed of .005 in/sand dwells for 30 seconds. Stage 5 - crimp head closes to .200 in at aspeed of .005 in/s and dwells for 30 seconds. Stage 6 - crimp headcloses to .160 in at a speed of .005 in/s and dwells for 30 seconds.Activate pressurization mode of crimping station to inflate a supportballoon with 50 psi to align any misaligned struts between Stage 3 andStage 5. After dwelling for 30 seconds the crimp head opens, remove thescaffold/support balloon from the crimp head. Remove partially crimpedscaffold and place it on the balloon of the balloon catheter (“FGballoon catheter”). Insert this assembly back into the center of thecrimp head. Reactivate the crimper. Stage 7 - crimp head closes to .160in at a speed of .25 in/s and dwells for 30 seconds. Stage 8 - crimphead closes to .130 in at a speed of .005 in/s and dwells for 50seconds. Activate pressurization mode to inflate the FG balloon catheter50 psi to create pillowing effect to improve scaffold retention anddwell for 50 seconds. Deactivate pressurization mode after 50 secondshave elapsed. Stage 9 - crimp head closes to .074 in at a speed of .005in/s and dwells for 150 seconds.

TABLE 3B crimping process for the “V59” scaffold Stage 1 - crimp headcloses to .3543 in at a speed of .3 inches per second (in/s) thenimmediately go to Stage 2. Stage 2 - crimp head closes to .270 in at aspeed of .005 in/s and dwells for 30 seconds. Stage 3 - crimp headcloses to .210 in at a speed of .005 in/s and dwells for 30 seconds.Stage 4 - crimp head closes to .160 in at a speed of .005 in/s anddwells for 30 seconds. Stage 5 - crimp head closes to .130 in at a speedof .005 in/s and dwells for 30 seconds. Activate pressurization mode ofcrimping station to inflate a support balloon with 50 psi to helpre-align any misaligned struts created during Stages 2 and 4. After the30 second dwell the crimp head opens, remove the scaffold/supportballoon from the crimp head. Remove partially crimped scaffold and placeit on the balloon of the balloon catheter (“FG balloon catheter”).Insert this assembly back into the center of the crimp head. Reactivatethe crimper. Stage 6 - crimp head closes to .140 in at a speed of .05in/s and dwells for 5 seconds. Stage 7 - crimp head closes to .130 in ata speed of .005 in/s and dwells for 30 seconds. Stage 8 - crimp headcloses to .100 in at a speed of .005 in/s and dwells for 30 seconds.Activate pressurization mode to inflate the FG balloon catheter to 50psi for dwell period. Stage 9 - crimp head closes to .0625 in at a speedof .005 in/s and dwells for 170 seconds.

Crimp processes similar to TABLES 3A and 3B are provided for V23 inparagraphs [0072] through [0092] in U.S. application Ser. No.12/861,719. In one embodiment, the crimping process may include the stepof maintaining an internal, elevated pressure (above ambient) or slowlybleeding a balloon-pressure valve while the crimping jaws are moving thescaffold diameter from a first to second diameter.

After Stage 9 the scaffold is immediately placed within a restrainingsheath to prevent recoil in the scaffold structure following crimping.The sheath is not intended to stay with the Finished Goods (FG) ballooncatheter when it is later implanted within a patient. Rather, the sheathis removed before the scaffold-balloon is introduced into the patient'saffected vessel lesion.

Examples of sheaths suited for this purpose are described in U.S.application Ser. No. 12/916,349. In the examples given here, variousslits, cuts or weakened areas may be pre-formed in the sheath tofacilitate a tearing away or removal of the sheath from the scaffold bya health professional, without dislodging the scaffold from the balloon.The removable sheath can have weakened areas preferably designed so thatit can be easily removed without applying an excessive pulling force onthe crimping scaffold.

Referring to FIGS. 12A-12B, a sheath 50 includes weakened areas 51 and53 arranged so that a medical professional may remove the sheath bytearing the sheath manually along lines of weaknesses in the sheath.Referring to FIG. 12A, when disposed over the crimped scaffold 22 of thescaffold-catheter assembly 18, the sheath 50 may have a proximal end 52and distal end 54 located, respectively, near the proximal and distalballoon markers 18 a, 18 b of the balloon 22 (FIG. 12A). At the distalend 54 there as a pair of opposed v-shaped cuts defining upper and lowerpull flaps 55 and 56. The dashed lines 53 indicate the intended tearline when the opposed flaps 55, 56 are manually pulled apart to initiatetearing along the line 53. In one embodiment, the tear lines 53 may alsocorrespond to pre-formed slits, having a depth about half the sheath 50wall thickness and over the length of the sheath 50. In one embodimentthe sheath 50 distal end 54 may extend beyond the scaffold distal end 14a of a distal portion 14 of the scaffold-catheter assembly 18. Thedistal end flaps 55, 56 may be folded up or over the sheath 50 to makethe flaps 55, 56 more easy to grip and pull up and down respectively, soas to avoid longitudinally applied forces on the scaffold surface as thesheath 50 is torn away.

Referring to FIG. 3, the pattern 200 includes longitudinally-spacedrings 212 formed by struts 230. A ring 212 is connected to an adjacentring by several links 234, each of which extends parallel to axis A-A.In this first embodiment of a scaffold pattern (pattern 200) four links234 connect the interior ring 212, which refers to a ring having a ringto its left and right in FIG. 3, to each of the two adjacent rings.Thus, ring 212 b is connected by four links 234 to ring 212 c and fourlinks 234 to ring 212 a. Ring 212 d is an end ring connected to only thering to its left in FIG. 3.

A ring 212 is formed by struts 230 connected at crowns 207, 209 and 210.A link 234 is joined with struts 230 at a crown 209 (W-crown) and at acrown 210 (Y-crown). A crown 207 (free-crown) does not have a link 234connected to it. Preferably the struts 230 that extend from a crown 207,209 and 210 at a constant angle from the crown center, i.e., the rings212 are approximately zig-zag in shape, as opposed to sinusoidal forpattern 200, although in other embodiments a ring with curved struts iscontemplated. As such, in this embodiment a ring 212 height, which isthe longitudinal distance between adjacent crowns 207 and 209/210 may bederived from the lengths of the two struts 230 connecting at the crownand a crown angle e. In some embodiments the angle e at different crownswill vary, depending on whether a link 234 is connected to a free orunconnected crown, W-crown or Y-crown.

The zig-zag variation of the rings 212 occurs primarily about thecircumference of the scaffold (i.e., along direction B-B in FIG. 3). Thestruts 212 centroidal axes lie primarily at about the same radialdistance from the scaffold's longitudinal axis. Ideally, substantiallyall relative movement among struts forming rings also occurs axially,but not radially, during crimping and deployment. Although, as explainedin greater detail, below, polymer scaffolds often times do not deform inthis manner due to misalignments and/or uneven radial loads beingapplied.

The rings 212 are capable of being collapsed to a smaller diameterduring crimping and expanded to a larger diameter during deployment in avessel. According to one aspect of the disclosure, the pre-crimpdiameter (e.g., the diameter of the axially and radially expanded tubefrom which the scaffold is cut) is always greater than a maximumexpanded scaffold diameter that the delivery balloon can, or is capableof producing when inflated. According to one embodiment, a pre-crimpdiameter is greater than the scaffold expanded diameter, even when thedelivery balloon is hyper-inflated, or inflated beyond its maximum usediameter for the balloon-catheter.

Pattern 200 includes four links 237 (two at each end, only one end shownin FIG. 3) having structure formed to receive a radiopaque material ineach of a pair of transversely-spaced holes formed by the link 237.These links are constructed in such a manner as to avoid interferingwith the folding of struts over the link during crimping, which, asexplained in greater detail below, is necessary for a scaffold capableof being crimped to a diameter of about at most Dmin or for a scaffoldthat when crimped has virtually no space available for a radiopaquemarker-holding structure.

A second embodiment of a scaffold structure has the pattern 300illustrated in FIG. 5. Like the pattern 200, the pattern 300 includeslongitudinally-spaced rings 312 formed by struts 330. A ring 312 isconnected to an adjacent ring by several links 334, each of whichextends parallel to axis A-A. The description of the structureassociated with rings 212, struts 230, links 234, and crowns 207, 209,210 in connection with FIG. 3, above, also applies to the respectiverings 312, struts 330, links 334 and crowns 307, 309 and 310 of thesecond embodiment, except that in the second embodiment there are onlythree struts 334 connecting each adjacent pair of rings, rather thanfour. Thus, in the second embodiment the ring 312 b is connected to thering 312 c by only three links 234 and to the ring 312 a by only threelinks 334. A link formed to receive a radiopaque marker, similar to link237, may be included between 312 c and ring 312 d.

FIGS. 6A and 6B depict aspects of the repeating pattern of closed cellelements associated with each of the patterns 300 and 200, respectively.FIG. 6A shows the portion of pattern 300 bounded by the phantom box VAand FIG. 6B shows the portion of pattern 200 bounded by the phantom boxVB. Therein are shown cell 304 and cell 204, respectively. In FIGS. 6A,6B the vertical axis reference is indicated by the axis B-B and thelongitudinal axis A-A. There are four cells 204 formed by each pair ofrings 212 in pattern 200, e.g., four cells 204 are formed by rings 212 band 212 c and the links 234 connecting this ring pair, another fourcells 204 are formed by rings 212 a and 212 b and the links connectingthis ring pair, etc. In contrast, there are three cells 304 formed by aring pair and their connecting links in pattern 300.

Referring to FIG. 6A, the space 336 and 336 a of cell 304 is bounded bythe longitudinally spaced rings 312 b and 312 c portions shown, and thecircumferentially spaced and parallel links 334 a and 334 c connectingrings 312 b and 312 c. Links 334 b and 334 d connect the cell 304 to theright and left adjacent ring in FIG. 4, respectively. Link 334 bconnects to cell 304 at a W-crown 309. Link 334 d connects to cell 304at a Y-crown 310. A “Y-crown” refers to a crown where the angleextending between a strut 330 and the link 336 at the crown 310 is anobtuse angle (greater than 90 degrees). A “W-crown” refers to a crownwhere the angle extending between a strut 330 and the link 334 at thecrown 309 is an acute angle (less than 90 degrees). The same definitionsfor Y-crown and W-crown also apply to the cell 204. There are eightconnected or free crowns 307 for cell 304, which may be understood aseight crowns devoid of a link 334 connected at the crown. There are oneor three free crowns between a Y-crown and W-crown for the cell 304.

Additional aspects of the cell 304 of FIG. 6A include angles for therespective crowns 307, 309 and 310. Those angles, which are in generalnot equal to each other (see e.g., FIG. 7A for the “V2” and “V23”embodiments of scaffold having the pattern 300), are indentified in FIG.5A as angles 366, 367 and 368, respectively associated with crowns 307,309 and 310. For the scaffold having the pattern 300 the struts 330 havestrut widths 361 and strut lengths 364, the crowns 307, 309, 310 havecrown widths 362, and the links 334 have link widths 363. Each of therings 312 has a ring height 365. The radii at the crowns are, ingeneral, not equal to each other. The radii of the crowns are identifiedin FIG. 6A as radii 369, 370, 371, 372, 373 and 374.

Cell 304 may be thought of as a W-V closed cell element. The “V” portionrefers to the shaded area 336 a that resembles the letter “V” in FIG.7A. The remaining un-shaded portion 336, i.e., the “W” portion,resembles the letter “W”.

Referring to FIG. 6B, the space 236 of cell 204 is bounded by theportions of longitudinally spaced rings 212 b and 212 c as shown, andthe circumferentially spaced and parallel links 234 a and 234 cconnecting these rings. Links 234 b and 234 d connect the cell 204 tothe right and left adjacent rings in FIG. 3, respectively. Link 234 bconnects to cell 236 at a W-crown 209. Link 234 d connects to cell 236at a Y-crown 210. There are four crowns 207 for cell 204, which may beunderstood as four crowns devoid of a link 234 connected at the crown.There is only one free crown between each Y-crown and W-crown for thecell 204.

Additional aspects of the cell 204 of FIG. 6B include angles for therespective crowns 207, 209 and 210. Those angles, which are in generalnot equal to each other (see e.g., FIG. 7B for the “V59” embodiment of ascaffold having the pattern 200), are indentified in FIG. 5B as angles267, 269 and 268, respectively associated with crowns 207, 209 and 210.For the scaffold having the pattern 200 the struts 230 have strut widths261 and strut lengths 266, the crowns 207, 209, 210 have crown widths270, and the links 234 have link widths 261. Each of the rings 212 has aring height 265. The radii at the crowns are, in general, not equal toeach other. The radii of the crowns are identified in FIG. 6A as innerradii 262 and outer radii 263.

Cell 204 may be thought of as a W closed cell element. The space 236bounded by the cell 204 resembles the letter “W”.

Comparing FIG. 6A to FIG. 6B one can appreciate that the W cell 204 issymmetric about the axes B-B and A-A whereas the W-V cell 304 isasymmetric about both of these axes. The W cell 204 is characterized ashaving no more than one crown 207 between links 234. Thus, a Y-crowncrown or W-crown is always between each crown 207 for each closed cellof pattern 200. In this sense, pattern 200 may be understood as havingrepeating closed cell patterns, each having no more than one crown thatis not supported by a link 234. In contrast, the W-V cell 304 has threeunsupported crowns 307 between a W-crown and a Y-crown. As can beappreciated from FIG. 6A, there are three unsupported crowns 307 to theleft of link 334 d and three unsupported crowns 307 to the right of link334 b.

The mechanical behavior of a scaffold having a pattern 200 verses 300differs in the following ways. These differences, along with others tobe discussed later, have been observed in comparisons between thescaffold of FIGS. 7A-7B, which include in-vivo testing. In certainregards, these tests demonstrated mechanical aspects of scaffoldsaccording to the invention that were both unexpected and contrary toconventional wisdom, such as when the conventional wisdom originatedfrom state of the art metallic stents, or coronary scaffolds. For aparticular design choice, whether driven by a clinical, productionyield, and/or delivery profile requirement, therefore, the followingcharacteristics should be kept in mind.

In general, a polymer scaffold that is crush-recoverable, possesses adesired radial stiffness and radial yield strength, fracture toughnessand is capable of being crimped down to a target delivery diameter,e.g., at least about D_(min), balances the three competing designattributes of radial yield strength and radial stiffness versestoughness, in-vivo performance verses compactness for delivery to avessel site, and crush recovery verses radial yield strength and radialstiffness.

In-vivo performance verses compactness for delivery to the vessel siterefers to the ability to crimp the scaffold down to the deliverydiameter. The ring struts 230 connecting crowns to form the W-cell 204are more restrained from rotating about an axis tangent to the abluminalsurface (axis A-A). In the case of the W-V cell the V portion, the crownmay tend to twist about the axis A-A under particular configurations dueto the reduced number of connecting links 336. The ring portions can ineffect “flip”, which means rotate or deflects out-of-plane as a resultof buckling (please note: “out-of-plane” refers to deflections outsideof the arcuate, cylindrical-like surface of the scaffold; referring toFIG. 5A “out-of-plane” means a strut that deflects normal to the surfaceof this figure). When there is a link 234 at each of a crown or valleyas in FIG. 5B, any tendency for the crown to buckle or flip is reducedbecause the ring struts are more restrained by the link 236.Essentially, the link serves to balance the load across a ring moreevenly.

The “flipping” phenomenon for a scaffold constructed according topattern 300 has been observed during crimping, as explained andillustrated in greater detail in U.S. application Ser. No. 12/861,719.The W-V cell 304 is devoid of a nearby link 334 at a crown 307 torestrain excessive twisting of the adjacent crown or valley. In essence,when there are two crowns 307 between a link 334 the restraintpreventing flipping or buckling of the V portion of the ring depends onthe buckling strength of the individual ring strut 330, i.e., the yieldstrength and stiffness of the polymer strut in torsion. When there is alink 234 connected to each adjacent crown/valley (FIG. 5B), however, outof plane deflections at the crown 207 is restrained more, due to thebending stiffness added by the connected link 234, which restrainstwisting at the adjacent crown 207.

A scaffold according to pattern 200 is correspondingly stiffer than asimilarly constructed scaffold according to pattern 300. The scaffoldaccording to pattern 200 will be stiffer both axially and inlongitudinal bending, since there are more links 236 used. Increasedstiffness may not, however, be desirable. Greater stiffness can producegreater crack formation over a less stiff scaffold. For example, thestiffness added by the additional links can induce more stress on ringsinterconnected by the additional links 234, especially when the scaffoldis subjected to a combined bending (rings moving relative to each other)and radial compression and/or pinching (crushing). The presence of thelink 234 introduces an additional load path into a ring, in addition tomaking the ring more stiff.

In-vivo requirements can favor a scaffold according to pattern 200, buta scaffold according to pattern 300 may be more easily crimped down tothe delivery diameter. Other factors also affect the ability to crimp ascaffold. According to the disclosure, it was found that crown anglesless than about 115 degrees for the pre-crimp scaffold can produce lessfracture and related deployment problems (e.g., uneven folding/unfoldingof ring struts) than scaffold with higher crown angles (relative to theinflated diameter, in one case 6.5 mm). The scaffold is crimped to aballoon that can be inflated up to about 7.4 mm. Thus, when the balloonis hyper-inflated the scaffold attains about up to about a 7 mm inflateddiameter. For a balloon catheter-scaffold assembly according to thedisclosure the largest inflated diameter for the balloon is less than orequal to the scaffold diameter before crimping. As mentioned above, itis preferred that the maximum inflated diameter for the scaffold is lessthan the scaffold diameter before crimping.

During the course of designing a crush recoverable polymer scaffoldhaving a desired crimped profile, it was found that when forming thescaffold at the 8 mm diameter it was difficult to crimp the scaffold toa desired crimped profile, e.g., to crimp the scaffold from the 8 mmdiameter to about 2 mm profile, for two reasons. First, by imposing the350-400% diameter reduction requirement, the polymer material was moresusceptible to crack formation and propagation, simply due to strainlevels experienced by the scaffold when subjected to this extensivediameter reduction. This concern was addressed by adjusting stiffness,e.g., reducing the strut angle, wall thickness and/or number of crowns.Additionally, the process steps used to form the tube (FIG. 1) improvesthe scaffold's resistance to crack formation and propagation, asexplained earlier.

Second, even when the scaffold dimensions were adjusted to limit crackformation, there was the problem of limited space for scaffold withinthe crimped profile. Due to the mass of material associated with thecrimped scaffold, the available space for compression of the rings tothe desired crimped profile was not achievable without creatingunacceptable yield stresses or fracture. Thus, even when a 350-400%diameter reduction was achievable without crack or deployment problems,the scaffold pattern would not allow further reduction without exceedingthe range of articulation that the scaffold design would allow.

According to another aspect of the disclosure, there are modified crowndesigns for a scaffold intended to improve the fracture toughness and/orreduce the delivery diameter of the scaffold. It was discovered that adesign change to an existing scaffold pattern that would overcome alimitation on reduced profile, and which could be implemented using abrittle polymer like PLLA of PLGA, was a significant reduction in thesize of the inner radius of the crown or valley bridging the struts thatform the crown/valley.

FIGS. 8A and 8B illustrate a pair of struts 420, 422 near a crown 410.In the pre-crimp state, the struts 420, 422 are separated by the crownangle φ and the crown is formed with an inner radius r_(a). This is atypical design for a crown. The inner radius is selected to avoid stressconcentrations at the crown. As the art has taught when there is adramatic change in geometry at a hinge point, such as a crown, there isa greater likelihood cracks or yielding will form at the hinge point(thereby affecting radial yield strength) since the moment of inertia inbending across the crown is discontinuous.

In the case of a metal stent, the angle φ before crimping is less thanthe angle when the stent is deployed. By forming the stent with thereduced diameter, the stent may be more easily crimped to a smallprofile. Due to the presence of the inner radius, the angle φ is capableof being exceeded at deployment without loss of radial stiffness. Ifthis radius is too small, however, and the strut angle at deploymentexceeds the angle before crimping, i.e., φ, there is a greater chance ofyielding or other problems to develop due to stress concentrations atthe inner radius. Due to the ductility and resiliency of metal, stentsmade from metal may also be crimped down further than shown in FIG. 8B.The struts 420, 422 may touch each other, i.e., S is less than 2×r_(a),and yet the stent can still recover and maintain its radial stiffnessdespite the over crimped condition.

For polymer scaffold, however, it has been found that the distance S(FIG. 8B) should not generally be smaller than allowed for the radiusr_(a), i.e., S greater than or equal to 2 r_(a). For a polymer scaffold,if the struts 420, 422 are brought closer to each other, i.e., S becomesless than 2×r_(a), the brittleness of the material can likely result infracture problems when the scaffold is deployed. The scaffold may nottherefore be able to maintain its radial stiffness if crimped beyond theallowable distance for the radius. The scanning electron microscope(SEM) photographs included as FIGS. 9A, 9F and 9G show fractures atcrowns when the distance S in FIG. 7B is less than 2×r_(a). As can beseen in these photographs, there is significant material failure in a Wcrown, free crown and Y crown.

With the objective of decreasing the distance S between struts 420, 422(FIG. 8B) the inventors decided to reduce down the radius r_(a) as smallas possible, despite the advice offered by the art. It was discovered,to their surprise, that the scaffold was able to recover from thecrimped condition to the expanded condition without significant,noticeable, reoccurring or prohibitive loss in radial yield strength.The SEMs provided as FIGS. 9B, 9C and 9D show crowns/valleys havingreduced radii after being crimped, then expanded by the balloon. Inthese examples the crown inner radii were made as small as the cuttingtool (a green light pico-second laser, described above) was able toproduce. As can be seen by comparing FIGS. 9A, 9F and 9G with FIGS. 9B,9C and 9D the scaffold having reduced radii produced some voids butthere is no crack propagation. Structural integrity was maintained. Thedeployed scaffold in these photos maintained good radial stiffness.

FIGS. 8C and 8D illustrate embodiments of a crown formation thatproduced these unexpected results. An example of a W cell having areduced radii type of crown formation just described is illustrated inFIG. 6B and 7B. The radius r_(b) is about 0.00025 inches, whichcorresponds to the smallest radius that could be formed by the laser.The 0.00025 inch radius is not contemplated as a target radius or limiton the radius size, although it has produced the desired result for thisembodiment. Rather, it is contemplated that the radius may be as closeto zero as possible to achieve a reduced profile size. The radius,therefore, in the embodiments can be about 0.00025 (depending on thecutting tool), greater than this radius, or less than this radius topractice the invention in accordance with the disclosure, as will beappreciated by one of ordinary skill in the art. For instance, it iscontemplated that the radii may be selected to reduce down the crimpedsize as desired.

An inner radius at about zero, for purposes of the disclosure, means theminimum radius possible for the tool that forms the crown structure. Aninner radius in accordance with some embodiments means the radius thatallows the distance S to reduce to about zero, i.e., struts are adjacentand/or touch each other as shown in FIG. 8D (S′ is about, or zero).

Without wishing to be tied to a particular theory for how the scaffoldaccording to the invention is capable of being reduced down to thetheoretical minimum diameter and then expanded without loss of strength,it is believed that the selection of starting diameter being greaterthan the inflated diameter played a role in the favorable outcome. Incontrast to the previous example where a metal stent is formed from adiameter less than its inflated diameter, which smaller diameter may beselected to facilitate a smaller crimped profile, a polymer scaffoldaccording to preferred embodiments is formed from a starting diametergreater than the maximum inflated diameter for the ballooncatheter-scaffold assembly (a larger starting diameter may be preferredto reduce acute recoil, as explained below, and/or to enhance radialyield strength characteristics in the deployed state as explainedearlier in the tube processing steps for the tube of FIG. 1). As such,the strut angle pre-crimp is preferably greater than the maximumcrown/strut angle when the scaffold is deployed. Stated differently, thecrown angle in FIG. 8C (pre-crimp angle) is never exceeded when theballoon expands the scaffold from the crimped to deployed state. Thischaracteristic of the crush recoverable polymer scaffold, i.e.,pre-crimp crown angle greater than the deployed crown angle, is believedto provide clues as to how the polymer scaffold in the SEM photographswas able to retain radial yield strength when a minimum inner radius wasused for the crown formation, contrary to the prior art. Compression,but not expansion of the scaffold when loaded by the vessel, it isbelieved, will not induce further weakening, despite the presence ofvoids. When the crown experiences only a compressive deformationrelative to its pre-crimp shape (FIG. 8C), the potentially weakened areanear the inner radius is subjected to only compressive stresses, whichdo not tend to tear the crown apart, i.e., induce crack propagation.

Crimping of the scaffold, as detailed in U.S. application Ser. No.12/861,719, includes heating the polymer material to a temperature lessthen, but near to the glass transition temperature of the polymer. Inone embodiment the temperature of the scaffold during crimping is raisedto about 5 to 10 degrees below the glass transition temperature forPLLA. When crimped to the final, crimped diameter, the crimping jaws areheld at the final crimp diameter for final dwell period. This method forcrimping a polymer scaffold having crush recovery is advantageous toreduce recoil when the crimp jaws are released. Another, unexpectedoutcome, however, was found relating to the reduced inner radius aspectof the disclosure. It was found that during the dwell period the polymerscaffold crimped profile could be reduced to a profile less than thetheoretical minimum profile.

From the example given earlier for the scaffold of FIG. 7B, the valuefor Dmin is 0.1048 in or 2.662 mm. When crimping this scaffold accordingto the crimping procedure summarized above and described in U.S.application Ser. No. 12/861,719, it was found that the scaffold could bereduced down to a crimped profile of 0.079 in or 2.0066 mm. Hence, thecrimped profile was less than Dmin for this scaffold. With this profilea protective sheath of 0.085 in OD could be placed over the scaffold.When a drug coating was disposed over the scaffold, the profile of thescaffold with sheath was 0.092 in. For this scaffold the range of radialyield strength was 0.45-0.65 N/mm, range of radial stiffness was1.00-1.20 N/mm and the crush recoverability was about 90% (50% crush).

It is believed that a reduced profile less than Dmin was achieved due toa compression of the material during the dwell period. Essentially, thepressure imposed by the crimping jaws during the dwell period at theraised temperature caused the struts forming the ring to be squeezedtogether to further reduced the crimped scaffold profile. According tothese embodiments, the crimped scaffold having a profile less than itstheoretical minimum profile was successfully deployed and tested invivo. This scaffold possessed the desired radial stiffness properties,in addition to the desired crush recovery of above about 90% following a50% reduction in diameter.

In another aspect of this disclosure, the strut and crown formation fora crush recoverable polymer scaffold is formed to take the shapedepicted in FIG. 8E, for purposes of achieving a crimped profile lessthan the crimped profile for the scaffold having the crown formationshown in FIG. 8A. According to these embodiments, the crown is formedwith a radius r_(c) as shown. When this scaffold is crimped, the strutsmay be brought close together so that the distance separating them isnear zero (S″ is about, or zero). In contrast to the embodiments of FIG.8C, the radius r_(c) is made some finite or larger radii than by forminga hole or enlarged area between the ends of the struts and crown. Thethickness at the crown, t_(c)′ forming the inner radius along its innersurface may be less than the strut width (in the example of FIG. 8C thecrown thickness may be larger than the strut width). This can allow alarger inner radius to be used at the crown without increasing thecrimped profile.

In these embodiments, a scaffold having the crown formation depicted inFIGS. 8E-8F is referred to as a “key-hole” crown formation. The namewill be understood without further clarification by reference to FIG.8F, which shows a key-hole slot or opening formed by the inner wallsurfaces. In the crimped profile, the struts near the crown may bebrought closer together while a hole or opening having radius r_(c) ismore or less maintained at the crown. The distance “S” is less thantwice the radius r_(c) for the “key-hole” crown formation.

Examples of scaffold embodying patterns 300 and 200 are provided inFIGS. 7A-7B (referred to as the V2 embodiment, which has a 0.008 inchwall thickness, V23 embodiments having 0.008 and 0.014 inch wallthickness and the V59 embodiment, which has a 0.011 inch wallthickness). Specific values for the various cell attributes of FIGS.6A-6B are provided.

The scaffold V59 (pattern 200) having a pre-crimp diameter of 8 mm iscapable of being crimped to a non-compliant balloon wherein the crimpedprofile is about 2 mm. The inflated diameter is about 6.5 mm in thisexample. The scaffold V2, V23 having pre-crimp diameters 7 and 9,respectively, are expanded to about 6.5 mm by a non-compliant balloon.The V2 and V23 scaffold are capable of being crimped to diameters ofabout 0.092 inches (2.3 mm).

According to the disclosure, it was found that the aspect ratio (AR) ofa strut of a scaffold may be between about 0.8 and 1.4, the AR of a linkmay be between about 0.4 and 0.9, or the AR of both a link and a strutmay between about 0.9 and 1.1, or about 1. Aspect ratio (AR) is definedas the ratio of width to thickness. Thus for a strut having a width of0.0116 and a wall thickness of 0.011 the AR is 1.05.

According to the disclosure, the radial yield strength of a balloonexpanded polymer scaffold having crush recoverability has a radial yieldstrength of greater than about 0.3 N/mm, or between about 0.32 and 0.68N/mm, and a radial stiffness of greater than about 0.5 N/mm or betweenabout 0.54 N/mm and 1.2 N/mm. According to the disclosure, acrush-recoverable scaffold has these ranges of stiffness and yieldstrength for a scaffold having a wall thickness of about 0.008 in to0.014 in and configured for being deployed by a 6.5 mm non-compliantballoon from about a 2 mm crimped profile, or deployed to a diameter ofbetween about 6.5 mm and 7 mm from about a 2 mm crossing profile on aballoon catheter.

A biodegradable polymer, such as PLLA (and polymers generally composedof carbon, hydrogen, oxygen, and nitrogen) is radiolucent with noradiopacity. It is desirable for a scaffold to be radiopaque, orfluoroscopically visible under x-rays, so that accurate placement withinthe vessel may be facilitated by real time visualization of the scaffoldbody, preferably the end rings. A cardiologist or interventionalradiologist typically will track a delivery catheter through thepatient's vasculature and precisely place the scaffold at the site of alesion using fluoroscopy or similar x-ray visualization procedures. Fora scaffold to be fluoroscopically visible it must be more absorptive ofx-rays than the surrounding tissue. Radiopaque materials in a scaffoldmay allow for its direct visualization. One way of including thesematerials with a biodegradable polymer scaffold is by attachingradiopaque markers to structural elements of the scaffold, such as byusing techniques discussed in U.S. application Ser. No. 11/325,973.However, in contrast to other stent or scaffold, a biodegradable,bioabsorbable, bioresorbable, or bioerodable, and peripherally implantedscaffold having crush recoverability according to the disclosure hasspecial requirements not adequately addressed in the known art.

There is the unmet need for maintaining a desired stiffness property inthe vicinity of the marker-holding material (marker structure) withoutincreasing the minimum crimped diameter, e.g., Dmin. The marker-holdingmaterial must not interfere with the extremely-limited space availablefor achieving the required crossing profile or delivery diameter for thecrimped scaffold on the delivery catheter, particularly in the case of ascaffold that has a diameter reduction of 300-400% or more when crimpedfrom the starting, pre-crimp diameter to the delivery diameter, and/orwhere the target delivery diameter is about at most a theoreticalminimum diameter (Dmin) for the scaffold. It has been found that inorder to be capable of achieving a desired delivery diameter, e.g.,300-400% or more diameter reduction during crimping, the marker material(when located on a link) should not interfere with the folding of thestruts forming rings of the scaffold. However, when addressing this needwithout consideration for the effect on radial stiffness, it was foundthat there was an unacceptable loss in stiffness in the vicinity of themarker structure.

Referring to FIGS. 10A and 10B there are shown portions of the scaffoldaccording to pattern 200. FIG. 10A shows the portion of the scaffoldwhere the link 237 holding a radiopaque material 500 (marker 500) islocated. FIG. 10B shows this same portion of the scaffold whenconfigured in a crimped configuration. The rings 212 b, 212 c, 212 d and212 f are shown in their compressed, folded or compact configuration ascrimped rings 212 b′, 212 c′, 212 d′ and 212 f′, respectively. So thateach of the rings 212 may have the same radial stiffness properties(ignoring link connections), the pair of markers 500 is preferablylocated on the link 237, as opposed to on a ring strut 230. In otherembodiments the marker 500 may be located on the ring 212 by makingsuitable accommodation in the ring structure.

As can be appreciated from FIG. 10B, in order to maintain the minimumdiameter, e.g., about at least the theoretical minimum crimped diameter(Dmin) for the crimped scaffold, the presence of marker structurepreferably has no effect on the distance between folded struts 230. Toachieve this result, the length of the link 237 may be increased,(L₂₃₇=L₁+L₂,) over the length L₁ of the other links 234 that do not havethe markers to carry (the length L₂ being about the length needed toaccommodate marker structure (depots 502 and the pair of markers 500),without interfering of limiting the folding of struts 230 as necessaryto achieve a 300-400% or more diameter reduction. Stents or scaffoldthat do not have a tight crimped diameter requirement or minimum spacebetween structural elements of a scaffold, by contrast, may have thelink connecting rings increased in size beneath the fold struts to holda marker 500, since there remains available space for marker structurein the crimped configuration.

The depots 502 may be formed when the scaffold is cut from the tube. Thedepots 502 provide a hole sized slightly smaller than a diameter of amarker 500 sphere, e.g., a platinum sphere, so that the sphere may beplaced in the hole and secured therein as a drug-polymer coating isapplied over the scaffold. The drug-polymer coating can serve as anadhesive or barrier retaining the marker 500 within the hole of a depot502.

In one aspect of the disclosure the diameter of a sphere forming themarker 500 necessary to achieve adequate illumination is less than thewall thickness (235, FIG. 4) of the polymer scaffold. As such, thesphere may be placed within the hole and then a coating applied over it.Since the sphere diameter is about equal to or less than the wallthickness 235 no reforming, or shaping of the sphere is necessary toachieve a flat profile. A process of applying the marker, therefore, issimplified.

When the length of a link having marker structure is increased tomaintain the minimum crimped diameter according to the embodiments ofFIG. 10, however, the combined radial stiffness properties of the nearbyrings is reduced since they are spaced further apart. To minimize thisloss in stiffness, particularly with respect to the end ring (which isinherently less stiff since it is connected to only one neighboringring), the marker structure is located between links 212 c and 212 f, asopposed to rings 212 d and 212 f. Additionally, the marker structure isarranged so that the marker pair 500 is placed in depots 502 a, 502 borientated along the vertical axis B-B as opposed to longitudinally(axis A-A). By placing the depots 502 a and 502 b along axis B-B thelength L₂ is preferably less than if the markers 500 were disposedlongitudinally, so that the undesirable loss in the combined radialstiffness of the adjacent rings 212 c, 212 f (resulting from theincreased length of link 237) and the end ring 212 d is minimal.

Design Process

As mentioned earlier, the problem may be stated in general terms asachieving the right balance among three competing design drivers: radialyield strength/stiffness verses toughness, in-vivo performance versescompactness for delivery to a vessel site, and crush recovery versesradial yield strength/stiffness.

Embodiments having patterns 200 or 300 were found to produce desiredresults with particular combinations of parameters disclosed herein, orreadily reproducible in light of the disclosure. It will be recognizedthere were no known predecessor balloon-expandable stents havingadequate crush recovery to use as a guide (indeed, the art haddiscouraged such a path of development for a peripheral stent). As such,various polymer scaffold combinations were fabricated based and thefollowing properties evaluated to understand the relationships bestsuited to achieve the following objectives:

Crush recoverability of the scaffold without sacrificing a desiredminimal radial stiffness and yield strength, recoil, deploy-ability andcrimping profile;

Acute recoil at deployment—the amount of diameter reduction within ½hour of deployment by the balloon;

Delivery/deployed profile—i.e., the amount the scaffold could be reducedin size during crimping while maintaining structural integrity;

In vitro radial yield strength and radial stiffness;

Crack formation/propagation/fracture when crimped and expanded by theballoon, or when implanted within a vessel and subjected to acombination of bending, axial crush and radial compressive loads;

Uniformity of deployment of scaffold rings when expanded by the balloon;and

Pinching/crushing stiffness.

These topics have been discussed earlier. The following providesadditional examples and conclusions on the behavior of a scaffoldaccording to the disclosure, so as to gain additional insight intoaspects of the disclosed embodiments.

A scaffold fabricated with a pattern similar to pattern 300 (FIG. 5)possessed a good amount of crush recoverability, however, thisscaffold's other properties were not ideal due to memory in the materialfollowing balloon expansion. The scaffold, which was initially formedfrom a 6.5 mm tube and deployed to about the same diameter, had acuterecoil problems—after deployment to 6.5 mm it recoiled to about a 5.8 mmdiameter. The scaffold also exhibited problems during deployment, suchas irregular expansion of scaffold rings.

One attempt at solving the design problem proceeded in the followingmanner. The scaffold's properties were altered to address stiffness,yield strength, structural integrity, deployment and recoil problemswhile maintaining the desired crush recoverability. Ultimately, ascaffold was designed (in accordance with the disclosure) having thedesired set of scaffold properties while maintaining good crush recoveryproperties after a 50% pinch deformation, which refers to the scaffold'sability to recover its outer diameter sufficiently, e.g., to about90-95%, following a crushing load that depresses the scaffold to aheight about equal to 50% of its un-deformed height.

The pinching stiffness (as opposed to the radial stiffness) is mostinfluenced or most sensitive to changes in the wall thickness of thescaffold. As the wall thickness increases, the pinching stiffnessincreases. Moreover, the crush recoverability of a scaffold is mostaffected by the stresses created at the regions that deflect mostoutward in response to the applied load. As explained below, as the wallthickness is increased, the crush recoverability decreases due to anincreased concentration of strain energy at the outwardly deflected endsof the scaffold. A design for a crush recoverable scaffold, therefore,must balance the wall thickness for increased pinching stiffness againstthe reduction in crush recoverability resulting from an increasedpinching stiffness. Similarly, although radial stiffness is lessaffected by changes in wall thickness (since loads are morepredominantly in-plane loading as opposed to out of plane duringpinching) when wall thickness is altered to affect crush recoverabilitythe radial stiffness must be taken into consideration. Radial stiffnesschanges when the wall thickness changes.

The diagrams drawn in FIGS. 11A, 11B and 11C are offered to assist withexplaining a relationship between wall thicknesses and crushrecoverability. FIG. 11A shows a cross-section of a scaffold in itsun-deformed (unloaded) state and deformed state when subjected to apinching load (drawn in phantom). The ends of the scaffold designated by“S” and “S” refer to regions with the highest strain energy, as one canappreciate by the high degree of curvature in these areas when thescaffold is under the pinching load. If the scaffold will not recover orhave reduction in recovery from the pinching load (F), it will bebecause in these regions the material has yielded which precludes orreduces recovery back to the pre-crush diameter. The equal and oppositecrushing forces in force F in FIG. 11A deflect the scaffold height fromits un-deformed height, i.e., the scaffold diameter, to a deformedheight as indicated by δ. The region of the scaffold that will containthe highest degree of strain energy when the crushing force F is beingapplied is near the axis of symmetry for the deformed shape, which isshown in phantom. In the following discussion, the load reaction ormaterial stress/strain state at the scaffold regions S and S′ will beexpressed in terms of the strain energy.

FIGS. 11B and 11C are simplified models of the loaded structure intendedto illustrate the effects on the strain energy in region S when thescaffold has different wall thickness. Essentially, the model attemptsto exploit the symmetry of the deformed shape in FIG. 11A to construct alinear stress-strain representation at region S in terms of a springhaving a spring constant K. Accordingly, the scaffold properties aremodeled as arcs 10/20 (½ of a hoop or ring) or half-cylinder shellssupported at the ends. The arc cannot displace downward (Y-direction)when the enforced displacement δ is applied, which is believedacceptable as a boundary condition due to the symmetry in FIG. 11A.Movement in the x-direction is restrained by the spring having springconstant K. The hemispherical arc 10 in FIG. 11C has a thickness t₁ andthe hemispherical arc 20 in FIG. 11B has a thickness of t₂>>t₁.

As the pinching load is applied in FIGS. 11B and 11C, the arcs 10 and 20are deformed (as shown in phantom). This is modeled by an enforceddisplacement of the arcs 10/20 at their center by about the amount delta(δ) as in FIG. 11A. The arc 10 deforms less than arc 20, however, interms of its curvature when the enforced displacement is applied,because its flexural rigidity is higher than arc 20. Since the curvatureis less changed in arc 10, more of the % strain energy resulting fromthe enforced displacement will be carried by the spring at the ends,where the spring force is restraining outward movement at S. For arc 20more % strain energy is carried in the arc, as the greater changes ofcurvature are intended to show, as opposed to the spring restrainingmovement at the ends.

Consequently, for a given applied force the % strain energy at the endswill be greater for arc 10, since the flexural rigidity of the arc 10 isgreater than the arc 20. This is depicted by the displacement of thespring (x₂>x₁). The % strain energy in the spring restraining arc 20(i.e., ½K(x₂)²/(total strain energy in arc 20)×100) is greater than the% strain energy in the arc 10 restraining spring (i.e., ½K(x₁)²/(totalstrain energy in arc 10)×100). From this example, therefore, one cangain a basic appreciation for the relationship between wall thicknessesand crush recoverability.

In a preferred embodiment it was found that for a 9 mm scaffoldpre-crimp diameter a wall thickness of between 0.008 in and 0.014 in, ormore narrowly 0.008 in and 0.011 in provided the desired pinchingstiffness while retaining 50% crush recoverability. More generally, itwas found that a ratio of pre-crimp or tube diameter to wall thicknessof between about 30 and 60, or between about 20 and 45 provided 50%crush recoverability while exhibiting a satisfactory pinching stiffnessand radial stiffness. And in some embodiments it was found that a ratioof pre-crimp or tube diameter to wall thickness of between about 25 and50, or between about 20 and 35 provided 50% crush recoverability whileexhibiting a satisfactory pinching stiffness and radial stiffness.

Wall thickness increases for increasing pinching stiffness may also belimited to maintain the desired crimped profile. As the wall thicknessis increased, the minimum profile of the crimped scaffold can increase.It was found, therefore, that a wall thickness may be limited both bythe adverse effects it can have on crush recoverability, as justexplained, as well as an undesired increase in crimped profile.

Provided below are results from various tests conducted on scaffolds andstents for purposes of measuring different mechanical properties andmaking comparisons between the properties of the stents and scaffolds.The stents used in the tests were the Cordis® S.M.A.R.T.® CONTROL® Iliacself-expanding stent (8×40 mm) (“Control stent”) the REMEDY Stent (6×40mm) by Igaki-Tamai (“Igaki-Tamai stent”), and the Omnilink Elite® stent(6×40 mm).

The data presented in Tables 4-5 (below) and Tables 4-6 in U.S.application Ser. No. 13/015,474 for the scaffolds V2, V23 and V59 arefor scaffolds having the properties listed in Tables 7A and 7B,respectively. The scaffolds were crimped to a delivery balloon, thenexpanded to their inflated diameter. The crimping process is similar tothat described at paragraphs [0071]-[0091] of U.S. application Ser. No.12/861,719.

The data presented in Tables 4-5 refer to scaffolds and stent propertiesafter they were expanded by their delivery balloons. For each of thetests reported in Tables 2-6, unless stated otherwise the statistic is amean value.

Table 4 presents data showing the percentage of crush recovery forvarious scaffold compared with other types of stents. The scaffolds andstents were crushed using a pair of opposed flat metal plates movedtogether to crush or pinch the stents and scaffold by the respectiveamounts shown in the tables. The test was conducted at 20 degreesCelsius.

Table 4 compares the crush-recoverability of the V2, V23 and V59scaffold to the Igaki-Tamai stent and Omnilink Elite® (6 mm outerdiameter and 40 mm length) balloon expandable stent. The crush periodwas brief (about 0 seconds).

TABLE 4 Approximate crush recovery using flat plate test at 20 Deg.Celsius (as percentage of starting diameter, measured 12 hours followingcrush) when crushed when crushed when crushed when crushed by an amountby an amount by an amount by an amount equal to 18% of equal to 33% ofequal to 50% of equal to 65% of Stent/ starting diameter startingdiameter starting diameter starting diameter scaffold type (18% crush)(33% crush) (50% crush) (65% crush) V23 (.008 in wall thickness) 99% 96%89% 79% V23 (.014 in wall thickness) 99% 93% 84% 73% V59 (.011 in wallthickness) 99% 96% 88% 80% Igaki-Tamai 99% 94% 88% 79% Omnilink Elite^((R)) 93% 80% 65% 49%

As can be seen in the results there is a dramatic difference between theV2, V23 and V59 crush recovery compared with the Omnilink Elite®coronary stent. The best results are achieved by the V23 (0.008 in wallthickness) and V59 scaffold when taking into consideration the radialyield strength and stiffness properties of these scaffold compared withthe Igaki-Tamai stent (see Table 5 of U.S. application Ser. No.13/015,474.

Table 5 compares the crush recovery behavior for a V23 scaffold with0.008 in wall thickness (FIG. 7A) following a 50% crush. The data showsthe percent crush recovery of the V23 scaffold following a brief(approximately 0 seconds), 1 minute and 5 minute crush at 50% of thestarting diameter.

TABLE 5 Approximate crush recovery of V23 (.008 in wall thickness) usingflat plate test at 20 Deg. Celsius (as percentage of starting diameter,measured 24 hours following crush) when crushed when crushed by anamount by an amount equal to 25% of equal to 50% of starting diameterstarting diameter Crush duration (25% crush) (50% crush) 0 second crush100%  99% 1 minute crush 99% 86% 5 minute crush 92% 83%

FIG. 13 (see U.S. application Ser. No. 13/015,474) shows the crushrecovery properties for the V59 scaffold when crushed to 50% of itsstarting diameter over a 24 hour period following removal of the flatplates. There are three plots shown corresponding to the recovery of thescaffold following a 0 second, 1 minute and 5 minute crush duration. Thescaffold diameter was measured at different time points up to 24 hoursafter the flat plates were withdrawn. As can be seen in these plots,most of the recovery occurs within about 5 minutes after the flat platesare withdrawn. It is contemplated, therefore, that an about 90% crushrecovery is possible for longer periods of crush, e.g., 10 minutes, ½hour or one hour, for scaffold constructed according to the disclosure.

When the pinching or crushing force is applied for only a brief period(as indicated by “0 sec hold time (50%)” in FIG. 13 of U.S. applicationSer. No. 13/015,474 tests indicate a recovery to about 95-99% of itsinitial diameter. When the force is held for 1 minute or 5 minute, testsindicate the recoverability is less. In the example of FIG. 13, it wasfound that the scaffold recovered to about 90% of its initial diameter.The 1 minute and 5 minute time periods being about the same suggeststhat any effects of the visco-elastic material succumbing to a plasticor irrecoverable strain when in a loaded state has mostly occurred.

In accordance with the disclosure, a crush-recoverable polymer scaffold(having adequate yield strength and stiffness properties, e.g., thestiffness and yield strength properties of the scaffold in Table 5 ofU.S. application Ser. No. 13/015,474, has a greater than about 90% crushrecoverability when crushed by an amount equal to about 33% of itsstarting diameter (33% crush), and a greater than about 80% crushrecoverability when crushed by an amount equal to about 50% of itsstarting diameter (50% crush) following an incidental crushing event(e.g., less than one minute); a crush-recoverable polymer scaffold has agreater than about 90% crush recoverability when crushed by an amountequal to about 25% of its starting diameter (25% crush), and a greaterthan about 80% crush recoverability when crushed by an amount equal toabout 50% of its starting diameter (50% crush) for longer duration crushperiods (e.g., between about 1 minute and five minutes, or longer thanabout 5 minutes).

An acute recoil problem was observed. In one example, a scaffold wasformed from a 7 mm deformed tube having a 0.008 in wall thickness. Whenthe scaffold was balloon deployed to 6.5 mm, the scaffold recoiled toabout 5.8 mm. To address this problem, the scaffold was formed fromlarger tubes of 8 mm, 9 mm and 10 mm. It was found that a largerpre-crimp diameter relative to the intended inflated diameter exhibitedmuch less recoil when deployed to 6.5 mm. It is believed that the memoryof the material, formed when the deformed tube was made, reduced theacute recoil.

A starting tube diameter of 10 mm, for example, for a scaffold having a7.4 mm inflated diameter should exhibit less recoil than, say, a 8 mmtube, however, this larger diameter size introduced other problems whichdiscouraged the use of a larger tube size. Due to the larger diameter itbecame difficult, if not infeasible to reduce the diameter duringcrimping to the desired crimped diameter of about 2 mm. Since there ismore material and a greater diameter reduction, there is less spaceavailable to reduce the diameter. As such, when the starting diameterexceeds a threshold, it becomes infeasible to maintain the desiredcrimped profile. It was found that a 9 mm tube size produced acceptableresults in that there was less recoil and a crimped profile of about 2mm could still be obtained.

An excessive starting diameter can introduce other problems duringdeployment. First, when the diameter reduction from starting diameter tocrimped diameter is too great, the local stresses in the scaffold hingeelements, crowns or troughs correspondingly increase. Since the polymermaterial tends to be brittle, the concern is with cracking or fractureof struts if stress levels are excessive. It was found that the diameter9 mm starting diameter scaffold (in combination with other scaffolddimensions) could be reduced down to 2 mm then expanded to the 7.4 mminflated diameter without excessive cracking or fracture.

As discussed earlier, unlike a metal stent, a design for a polymerscaffold must take into consideration its fracture toughness both duringcrimping and when implanted within a vessel. For a scaffold locatedwithin a peripheral artery the types of loading encountered are ingeneral more severe in terms of bending and axial loading than acoronary scaffold, in addition to the pinching or crush forcesexperienced by the scaffold, due to the scaffold's proximity to thesurface of the skin, and/or its location within or near an appendage ofthe body. See e.g. Nikanorov, Alexander, M. D. et al., Assessment ofself-expanding Nitinol stent deformation after chronic implantation intothe superficial femoral artery.

As is known in the art, a scaffold designed to have increased radialstiffness and yield strength properties does not, generally speaking,also exhibit the fracture toughness needed for maintaining structuralintegrity. The need to have a peripherally implanted polymer scaffoldwith adequate fracture toughness refers both to the need to sustainrelatively high degrees of strain in or between struts and links of thescaffold and to sustain repeated, cyclical loading events over a periodof time, which refers to fatigue failure.

The methods of manufacture, discussed earlier, of the tube from whichthe scaffold is formed are intended to increase the inherent fracturetoughness of the scaffold material. Additional measures may, however, beemployed to reduce instances of fracture or crack propagation within thescaffold by reducing the stiffness of the scaffold in the links, or byadding additional hinge points or crown elements to the ring.Alternatively or in addition, pre-designated fracture points can beformed into the scaffold to prevent fracture or cracks from propagatingin the more critical areas of the scaffold.

Cracking/fracture problems are also observed as a consequence ofirregular crimping and/or deployment of the scaffold. Irregulardeployment is problematic, not only from the viewpoint of the scaffoldnot being able to provide a uniform radial support for a vessel, butalso from the viewpoint of crack propagation, fracture and yielding ofstructure resulting in loss of yield strength and/or stiffness in vivo.Examples of irregular deployment include crowns being expanded beyondtheir design angles and in extreme cases, flipping or buckling of crownsduring deployment or crimping. These problems were observed duringcrimping process and during deployment, examples of which are describedin greater detail in U.S. application Ser. No. 12/861,719.

Pattern 300 may be susceptible to more of these types of problems thanpattern 200. The links of the pattern provide less support for the ringstruts forming the V segment of the W-V closed cell 304, as compared topattern 200. It is believed that the w-shaped closed cell 204 was morecapable of deploying without irregularities, such as flipping, due toits symmetry. The asymmetric loading inherent in the W-V cell 304 wasmore susceptible to buckling problems during crimping or deployment.These potential problems, however, should they arise, may be addressedby adopting modifications to the crimping process.

For example, a scaffold having a diameter of 7 mm and asymmetric closedcells (pattern 300) was crimped then deployed without any flipping ofstruts observed. A second scaffold of 9 mm diameter was then crimped toa balloon and deployed. This scaffold had the same pattern 300 as the 7mm scaffold. The strut or crown angle was increased by the ratio of thediameters, i.e., increased by a factor of 9/7, to compensate for thechange in radial stiffness resulting from the increased diameter. Whenthe 9 mm scaffold was crimped, however, flipping occurred in thescaffold struts (primarily in the V section of the W-V closed cell). Tocorrect this problem the W closed cell (pattern 200) was tested. Thismodification helped to reduce instances of flipped struts. Surprisingly,the same irregular crimping/deployment problems have not been observedfor the comparable metal stent having a W-V closed cell pattern. It wasconcluded, therefore, that the flipping problem (in particular) is aphenomenon unique to a polymer scaffold.

To avoid flipping phenomena, should it occur in a metal stent, one mightconsider simply adjusting the moment of inertia of a strut to preventout of plane (outside of the arcuate, abluminal surface) deflection of astrut. However, as noted earlier, the polymer material introducesdegrees of freedom or limitations that are not present with a metallicmaterial. In the case of minimizing undesired motion of a strut bymodifying bending inertia properties of the strut one needs to bemindful that polymer struts must, generally speaking, be thicker and/orwider than the equivalent metal strut. This means there is less spaceavailable between adjacent struts and already higher wall thicknessesthan the metal counterpart. This problem of space is further compoundedfor embodiments that form a polymer scaffold from a tube that is theformed at the deployed, or larger than deployed size for the scaffold.It is desirable to have the scaffold reduced in diameter during crimpingfor passage to the same vessel sites as in the case of the metal stent.Thus, the delivery profile for the crimped scaffold should be about thesame as the metal stent.

A metal balloon expandable stent may be cut from a tube that is betweenthe deployed and crimped diameters. As such, the spacing between strutsis greater and the stent is more easily compressed on the balloonbecause the stent pre-crimp has a diameter closer to the crimpeddiameter. A polymer scaffold, in contrast, may be cut from a diametertube equal to or greater than the deployed state. This means there ismore volume of material that must be packed into the delivery profilefor a polymer scaffold. A polymer scaffold, therefore, has morerestraints imposed on it, driven by the crimped profile and startingtube diameter, that limits design options on strut width or thickness.

A well known design requirement for a vessel supporting prosthesis,whether a stent or scaffold, is its ability to maintain a desired lumendiameter due to the inward radial forces of the lumen walls includingthe expected in vivo radial forces imparted by contractions of the bloodvessel. Referring to the examples in FIGS. 7A-7B, the radial stiffnessand radial yield strength of the scaffold is influenced by the width ofstruts, crown radii and angles, length of ring struts extending betweencrowns and valleys, the number of crowns and the wall thickness(thickness 235, FIG. 4) of the scaffold. The latter parameter (wallthickness) influences the pinching stiffness, as explained earlier.During the design process, therefore, this parameter was altered toaffect pinching stiffness and crush recoverability, although it also hasan effect on radial stiffness. In order to affect the radial stiffness,one or more of the foregoing parameters (crown angle, crown radius, ringstrut length, crown number, and strut width) may be varied to increaseor decrease the radial stiffness.

To take one example, when it was found that a 7 mm scaffold's recoilproblem could be overcome by increasing the starting tube diameter to 8mm, 9 mm or perhaps even 10 mm, an initial approximation to thecorresponding changes to the scaffold pattern dimensions involvedincreasing characteristics such as ring strut length, crown angle andlink by the ratio of the diameters, e.g., 8/7 when increasing OD from 7mm to 8 mm. However, this rough approximation was found to beinsufficient in retaining other desired properties, such as crushrecoverability. Thus, further refinements were needed.

The relationships between radial stiffness and above mentionedparameters are well known. However, the relationship of thesestiffness-altering parameters to crush recoverability of a balloonexpandable stent, much less a balloon expandable scaffold is not wellknown, if known at all in the existing art. Accordingly, the designprocess required the constant comparison or evaluation among radialstiffness, pinching stiffness and crush recoverability (assuming thechanges did not also introduce yield or fracture problems duringcrimping and deployment) when the stiffness parameters were altered todetermine whether these and related scaffold properties could beimproved upon without significant adverse effects to crushrecoverability.

Tables 4-6 and the accompanying text of U.S. application Ser. No.13/015,474 provide values for acute recoil, radial yield strength andstiffness and pinching stiffness for scaffold produced by the processesof the invention, which are considered part of this disclosure.

While particular embodiments of the present invention have been shownand described, it will be obvious to those skilled in the art thatchanges and modifications can be made without departing from thisinvention in its broader aspects. Therefore, the appended claims are toencompass within their scope all such changes and modifications as fallwithin the true spirit and scope of this invention.

What is claimed is:
 1. A method for making a medical device, comprising:forming a scaffold from a tube made from a polymer; crimping thescaffold to a balloon, wherein the scaffold is plastically deformed tohave a crimped state; and fitting a removable sheath over the scaffoldand balloon following crimping to limit recoil of the crimped scaffold;wherein the scaffold has an expanded diameter when plastically deformedfrom the crimped state by the balloon; and wherein the scaffold attainsgreater than 90% of the expanded diameter after being crushed by anamount equal to at least 33% of the expanded diameter.
 2. The method ofclaim 1, wherein the scaffold is heated to a temperature below a glasstransition temperature of the polymer when the scaffold is being crimpedto the balloon.
 3. The method of claim 1, wherein the polymer is poly(L-lactide) (PLLA), a polymer made from at least 80% L-lactide, a blockcopolymer with a PLLA block, or a copolymer of PLLA.
 4. The method ofclaim 1, wherein the scaffold has a pre-crimp diameter prior to crimpingand a wall thickness, and a ratio of the pre-crimp diameter to the wallthickness is between 20 and
 45. 5. The method of claim 1, wherein thescaffold is crimped to a balloon of a balloon-catheter having a distalend, wherein the sheath includes a distal end that extends beyond thecatheter distal end, and wherein the sheath includes flaps that arefolded to enable the flaps to be gripped and pulled apart when removingthe sheath from the scaffold.
 6. The method of claim 1, wherein thescaffold includes rings having crowns and the crowns form crown angles,wherein during crimping a ring is reduced in diameter by plasticdeformation resulting in an articulation of struts about crowns; andwherein prior to crimping the scaffold has an outer diameter of 8 to 10mm, the crown angles for the rings are between 90 and 115 degrees, andthe scaffold has a wall thickness of at least 0.008 in.
 7. The method ofclaim 1, wherein the tube has a crystallinity of between about 30% and40%.
 8. The method of claim 1, wherein the tube is made from a polymercomprising a copolymer of poly (L-lactide) (PLLA) and polycaprolactone(PCL).
 9. A method for making a medical device, comprising: forming ascaffold from a tube made from a polymer, the scaffold having a firstdiameter; crimping the scaffold to a balloon by plastic deformation ofthe scaffold, wherein the scaffold is crimped from the first diameter toa second diameter, thereby forming a crimped state of the scaffold, thefirst diameter is greater than a nominal inflation diameter for theballoon, and the scaffold has an expanded diameter when plasticallydeformed from the crimped state by the balloon; and wherein the scaffoldis capable of regaining more than 90% of the expanded diameter afterbeing crushed by an amount equal to at least 33% of the expandeddiameter.
 10. The method of claim 9, wherein the scaffold has an 8 mminner diameter that is the first diameter and the balloon has a nominalinflated diameter of 6.5 mm.
 11. The method of claim 9, wherein thescaffold has a network of interconnected elements including strutsjoined at crowns to form rings and links connecting the rings, theplastic deformation causes an articulation of struts about crowns, andwherein an aspect ratio (AR) of a width to a wall thickness of a strutor link is between 0.4 and 1.4.
 12. The method of claim 11, wherein theAR is between 0.4 and 0.9.
 13. The method of claim 11, wherein the AR isbetween 0.8 and 1.4.
 14. The method of claim 9, wherein the scaffold hasa ratio of the first diameter to a wall thickness of between 20 and 60.15. The method of claim 14, wherein the ratio of the first diameter tothe wall thickness is between 20 and
 45. 16. The method of claim 14,wherein the ratio of the first diameter to the wall thickness is between30 and
 60. 17. A method, comprising: forming a scaffold from a tube, thescaffold comprising rings, each ring having crowns and struts, crimpingthe scaffold to a balloon by plastic deformation of the scaffold andwhile the scaffold is heated to an elevated temperature below a glasstransition temperature (Tg) of the polymer, wherein the scaffold iscrimped from a first diameter to a second diameter, and the firstdiameter is greater than a nominal inflation diameter for the balloon,and the crimping further including the step of performing a dwell of thescaffold after reducing the scaffold to the second diameter, wherein thesecond diameter is less than a minimum crimped diameter (MCD) satisfyingeither of equations 1 or 2:MCD =(Σ Swi +Σ Crj +ΣLwk) *(π)⁻¹+2*WT  (equation 1)MCD =(Σ Swi +Σ Crj +ΣLwk) *(π)⁻¹  (equation 2) such that Σ Swi (i=1. . .n) is the sum of n ring struts having width Swi; Σ Crj (j=1. . . m) isthe sum of m crown inner radii having radii Crj (times 2); Σ Lwk (k=1. .. p) is the sum of p links having width Lwk; and WT is a ring wallthickness.
 18. The method of claim 17, wherein the tube is made from apolymer comprising a copolymer of poly (L-lactide) (PLLA) andpolycaprolactone (PCL).
 19. The method of claim 17, wherein the scaffoldhas a wall thickness and a ratio of the first diameter to the wallthickness is between 20 and
 45. 20. The method of claim 19, wherein anaspect ratio (AR) of a width of a strut to the wall thickness is between0.8 and 1.4.